A male high-school athlete was participating in a team sport when a player from the opposing team attempted a goal. The male athlete was the only obstacle between the opposing player and a winning goal. The high speed shot, taken from less than 10 feet away, impacted the male athlete directly in the groin. He immediately fell to his knees in pain. Thankfully, he was wearing an new athletic protector (known colloquially as a “jockstrap”), which should have prevented injury even at such close quarters. Dr. John Lloyd was retained to perform a biomechanical analysis athletic protector.
The athlete sat out the remainder of the game. Later that evening he became concerned as the swelling continued. The following day tests revealed that amputation of one of his testicles was medically necessary. As a young man, with his whole life ahead of him, the physical and emotional pain of losing a testicle was almost unbearable.
The young man had conducted his research before purchasing the new athletic protector. The packaging had promised comfort and protection. Why then did he sustain this life-changing injury?
Athletic protector biomechanics expert Dr. John Lloyd, was retained to evaluate a potential product liability case.
It was quickly discovered, interestingly, that there are no American Standards on the performance requirements of athletic protectors. Therefore, Dr. Lloyd devised a test method to evaluate exemplars of the subject jockstrap with comparison to models sold by other product manufacturers.
Balls were shot at various speeds from a pitching machine aimed at the athletic protectors affixed to a male mannequin. Each impact was recorded using a high-speed video camera, while Dr. Lloyd’s associate, standing behind the mannequin, measured the speed of each impact using a radar gun. A total of 70 tests were performed.
As the following high-speed video recording shows, the subject athletic protector deforms completely upon impact, providing the wearer with little, if any, protection from injury.
Several new design models also collapsed upon impact, while others cracked and broke
Fortunately, the old style jock strap with which many of us are familiar was among the few models that held up to impact and actually provided adequate protection.
Based on biomechanical analysis I concluded, to a reasonable degree of scientific certainty, that the subject athletic protector provides inadequate protection of the male genitalia from injury associated with impact from a moderate speed ball. This conclusion is based on evidence of extreme deformation of the jock strap upon direct impact from a ball.
Had the manufacturer evaluated their product under real-life conditions, as described herein, they would have learned that this product provides inadequate protection against injury to the male genitalia. Further, comparative testing of other available athletic protectors identified products that provide better protection.
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.
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 peer-reviewed article on Motorcycle Accident Reconstruction, which was originally published in the Journal of Forensic Biomechanics in January 2016.
Corresponding author: John D Lloyd, Research Director, BRAINS, Inc., 32824 Michigan Avenue San Antonio, Florida, 33576, USA, Tel: 813-624-8986; Fax: 352-588-0688; E-mail: drjohnlloyd@tampabay.rr.com
Abstract
In a motorcycle accident, the motorcycle and rider typically become independent, each following their own path to final rest. Consequently, the biomechanical analysis of a motorcycle accident reconstruction is complex. A biomechanical model to assess rider kinematics associated with motorcycle accidents is presented, which may be important to forensic scientists involved in the analysis of such events. This model can also be applied to other activities, including cycling, equestrian sports, skiing, skating, running, etc.
In a motorcycle accident reconstruction, it is first important to understand the mechanisms by which a rider may be ejected from their motorcycle and how drag factors affect the motorcycle and rider independently. Next we determine rider trajectory, taking into consideration rider anthropometry and posture, results from which are used to derive impact velocity as a function of linear and angular components. A case study is presented, demonstrating how the presented model can be applied to a collision involving a single motorcycle.
Introduction to Motorcycle Accident Reconstruction
Motorcycles are a luxury in the developed world, where they are used mostly for recreation. Whereas in developing countries, motorcycles are required for utilitarian purposes due to lower prices and greater fuel economy. It is estimated that in 2016 there will be more than 134 million motorcycles worldwide [1], 60-80% of which are in the Asia Pacific and Southern and Eastern Asia regions. In 2011 there were more than 8.2 million registered motorcycles in the United States [2], representing 3% of all US registered vehicles, with California, Florida and Texas leading the number of motorcycles per state [3].
3.1. Epidemiology of motorcycle accidents
In the United States motorcyclists travelled 18.5 billion miles in 2011, which represents only 0.6% of total vehicle miles travelled, yet motorcyclists accounted for 14% (4,612) of traffic fatalities and 4% (81,000) of all occupant injuries [2]. According to the U.S. National Highway Traffic Safety Administration (NHTSA), when compared with automobiles, per vehicle mile traveled, motorcyclists’ risk of a fatal crash is 35 times greater than that of a car occupant [4].
Two major epidemiologic studies into the causation of motorcycle accidents have been conducted in North America and Europe: the Hurt Report and the MAIDS report. The Hurt Report [5] showed that failure of motorists to detect and recognize motorcycles in traffic is the prevailing cause of motorcycle accidents. Seventy-five percent of accidents were found to involve a motorcycle and a passenger vehicle, while the remaining 25% of accidents were single motorcycle accidents. Two-thirds of motorcycle-car crashes occurred when the car driver failed to see the approaching motorcycle and violated the rider’s right-of-way. Findings of the Hurt study indicate that severity of motorcyclist injury increase with speed, alcohol consumption, motorcycle size and speed.
The MAIDS study (Motorcycle Accidents In Depth Study) [6] is the most recent epidemiologic study of accidents involving motorcycles, scooters and mopeds, which was conducted in 1999 across five European countries to investigate motorcycle accident exposure data. Key findings show that passenger cars were the most frequent collision partner (60%), where 69% of the drivers report that they did not see the motorcycle and the predominance of motorcycle accidents (54.3%) occurred at an intersection.
In the United States alone, it is estimated that the total direct costs associated with motorcycle crashes in 2010 was approximately $16 billion. However, the US Government Accountability Office (GAO) predicts that full costs of motorcycle crashes are likely considerably higher because some difficult-to-measure costs, such as longer-term medical costs, are not included [7].
Biomechanical Model
A new model is presented for the purpose of investigating motorcycle accident reconstruction biomechanics involving a lone motorcycle, which accounts for 25% of all motorcycle-related accidents according to the Hurt report [5]. This model is unique in that it incorporates measures of rider anthropometry (body measurements) and riding posture, which have a direct effect on trajectory and overall height of the vertical component of the impact.
The model presented herein may be applied not only to motorcycle accidents, but also to a wide range of activities, including cycling, equestrian sports, skiing, skating, running, etc.
Methods
It is first important to understand the mechanisms by which a rider may be ejected from their motorcycle and how drag factors affect the motorcycle and rider independently. Next we determine rider trajectory, results from which are used to derive impact velocity as a function of linear and angular components. Finally, characteristics of the impact surface are considered with respect to impact accelerations.
5.1. Rider ejection
There are a number of ways that a rider can be ejected from the bike in a lone motorcycle accident. Two common ways of ejection are the lowside (Figure 1A) and highside (Figure 1B) crash. A rider may also be ejected over the handlebars (Figure 1C).
Figure 1 – Rider Ejected from Motorcycle
The lowsider or lowside is a type of motorcycle crash usually occurring in a turn (Figure 1A). A lowside crash is caused when either the front or rear wheel slides out as a result of either too much braking in a corner, too much acceleration through or out of a corner, or too much speed carried into or through a corner for the available traction. A lowside crash may also be caused by unexpected slippery or loose material (such as oil, water, dirt or gravel) on the road surface.
A highsider or highside is a type of motorcycle accident characterized by sudden and violent rotation of the motorcycle about its longitudinal axis. This generally happens when the rear wheel loses traction, skids, and then suddenly regains traction, creating a large torque, ejecting the rider off the side of the motorcycle, oftentimes head-first (Figure 1B).
Highside and lowside accidents differ as follows: during a lowside the rear wheel slips laterally and continuously until the motorcycle falls onto the side facing the inside of the corner.Whereas during a highside crash the rear wheel slips laterally before suddenly regaining traction and flipping the motorcycle toward the outside of the corner (the higher side of the motorcycle). Highsides happen quickly and are very violent consequently injuries tend to be more severe in a high side crash, compared to a lowside crash.
Endo, short for “end over end,” occurs when the front end of a motorcycle stays fixed while the rear rotates up into the air, causing a rider to fly over the handlebars (Figure 1C).
5.2. Drag factors
Drag factors for motorcycles have been established based on motorcycle accident reconstruction and typically range from 0.2-1.0 [8], where 0.25 represents a motorcycle with a fairing [9], such as a sport motorcycle. Sport and sport touring motorcycles will likely slide further than a cruiser-style motorcycle, which have more external components that resist sliding, for which a drag factor of 0.5 is commonly adopted Table 1, below, presents drag factors for street motorcycles sliding on typical road surfaces [10-13].
Table 1: Drag Factors for Sliding Motorcycles
Drag factors for the rider are typically higher than those for motorcycles sliding on a dry asphalt or concrete roadway. An extensive series of motorcycle accident reconstruction tests were carried out by the West Midlands Police in the United Kingdom in which they calculated the drag factor value of a crash-test dummy sliding across an airfield surface. The resulting coefficient was found to vary between 0.57 and 0.85 for normal clothing [14].With similar drag factors for dry and wet roadway conditions [15]. For the purpose of accident reconstruction, a drag factor of 0.7 for a clothed individual sliding on a roadway is generally accepted.
Evidence from final rest positions of the motorcycle and rider can be used to establish whether the rider was involved in a lowside or highside motorcycle ejection. In a lowside crash the motorcycle will tend to slide further than the rider. Whereas in a highside crash, the rider is ejected from the motorcycle, traveling additional distance over ground in a similar direction to the motorcycle, prior to making contact with the ground and initiating the slide. Hence, in a highside crash, the final rest position of the rider may be beyond the final rest position of the motorcycle. Furthermore, in higher-energy ejection crashes the rider is more likely to both slide and tumble, resulting in a longer travel distance from location of ejection from the motorcycle, as well as additional injuries as evidenced by fractures, lacerations and contusions to various regions of the body.
5.3. Rider anthropometry
Anthropometry is the study of human body measurements. Rider anthropometry will directly affect fall height, since head center of mass (HCOM) and overall center of mass (RCOM) varies between individuals.
In a lowside crash, seated height of the center of mass (HCOM) of the rider’s head approximates vertical fall height. Whereas in a highside crash vertical fall height is a function of seated head CoM height (HCOM), plus additional height gained based on trajectory of the rider calculated with reference to overall center of mass of the rider (RCOM) (Figure 2).
Figure 2 – Fall Height Associated with Low side and High side Accidents
5.4. Rider center of mass
Rider center of mass height (RCOM) is located anatomically with respect to the second sacral vertebra (S2), which can be visually estimated as approximate in height to the omphalion (navel), and is measured vertically with respect to the road surface. The preferred method for determining RCOM is to measure seated height of the rider on the subject motorcycle (Figure 3). If the rider is not available due to injury or fatality, then an exemplar same-gender person of similar height and weight may be used. If the subject motorcycle is not available due to extent of damage, then an exemplar motorcycle should be obtained. With the motorcycle supported perpendicular to the road by an assistant (not on side stand or center stand), and rider’s hands on the handlebar grips and feet on the foot pegs, measure the vertical height from the ground to the motorcycle seat at the location of the ischial tuberosities (base of the pelvic bones at the seat surface). An anthropometer and spirit level should ideally be used for accuracy and measurements recorded in millimeters to maximize precision.
Figure 3 – Rider anthropometry
As an alternative method, rider seated height can be calculated by sourcing motorcycle seat height, from manufacturer specifications, from which a correction factor for suspension compression under mass of the rider is subtracted. Suspension compression, also known as sag, will vary by motorcycle type and mass of the rider. A general rule of thumb is that the front sag should be about 30-35% of travel, while the rear should be at about 25%, which equates to 30-40 mm at the front and 25-35 mm at the rear for most bikes [16]. Therefore, a reasonable correction factor for suspension sag is 30-35 mm.
For both methods, an adjustment must be added to the compressed seat height to determine RCOM. According to Pheasant [17], seated RCOM is equal to seat height plus 10% of total stature (standing height). This factor calculation is identical for both males and females (Table 2).
Table 2: Anthropometric Data
5.5. Head center of mass
Head center of mass (HCOM) height can also be measured directly using the method described earlier, with the rider seated on the subject motorcycle in the correct riding position. An anthropometer is used to measure the vertical height from the ground. If the rider and/or subject motorcycle is unavailable, a substitute individual of similar height and weight and exemplar motorcycle may be used. The canthus (outer corner of the eye) is used as an anatomical landmark reference, equal in height to the center of mass of the head (Figure 3).
Alternatively, seated head center of mass (HCOM) can be calculated as a function of stature (standing height). Utilizing data from the 1988 Anthropometric Survey of U.S. Personnel [18], HCOM is derived by multiplying stature by 45.2%. Similar to the RCOM calculation, HCOM must be corrected for posture by multiplying HCOM by the cosine of seated back angle (β), measured with respect to the vertical axis.
5.6. Trajectory of the rider
The trajectory is the path that a rider is thrown or vaulted under the action of gravity, neglecting all other forces, such as friction from air resistance, without additional propulsion (Figure 4) and is defined by Equation 1.
Equation 1 – Trajectory of the ejected rider (y):
Figure 4 – Trajectory of an Ejected Rider
The following standard mathematical formulae are used to determine specific components of trajectory that are pertinent to the kinematic analysis of a rider ejected from a motorcycle.
5.7. Distance travelled
In a motorcycle accident reconstruction it may be possible to establish the actual distance of ejected travel of the rider, based upon location of ejection, typically between the end of any tire skid marks and start of gouge marks on the roadway, and location of bodily impact with the ground, identified by helmet paint transfer and/or identification of clothing fibers or body tissue on the roadway consistent with the rider. If the speed of the motorcycle at ejection (νejection) is also known, then the distance travelled (d) can be computed using the formula below, taking into account any correction factors for relative change in road height from location of ejection to impact.
Equation 2 – Horizontal distance traveled (d):
However, the location of bodily impact is often difficult to identify, in which case reasonable assumptions may be made, including utilization of an estimated ejection angle (θ).
5.8. Maximum height
One of the most critical factors for determination of total impact velocity in a motorcycle accident reconstruction is the maximum height attained by an ejected rider, which is calculated according to
Equation 3 – Maximum height (h):
5.9. Ejection angle
The ejection angle (θ) is the angle at which a rider must be launched in order to travel a certain distance, given the initial velocity. Oftentimes, based on the final rest positions of the motorcycle and rider and in consideration of appropriate drag factors, it is possible to approximate rider ejection velocity. Air resistance is considered negligible, therefore angle and velocity at ejection are considered equal to the angle and velocity at impact.
Equation 4 – Ejection angle (θ):
5.10. Rider impact velocity
Total impact velocity is derived on the basis of its vertical, angular and travel velocity components.
5.11. Linear vertical impact velocity
Vertical impact velocity is computed as a function of seated head height, plus any additional height gained due to rider ejection from the motorcycle. The potential energy (P.E.) at any point will depend on the mass (m) at that point and its distance above the ground (h), multiplied by the gravitational acceleration constant (g) (Figure 5).
Figure 5 – Potential Energy of a Motorcyclist
The potential energy of the entire system is the integral of the energies of each finite mass element of the motorcycle plus rider over its height: . For simplification, we assume that the mass is evenly distributed over the system. Hence, P.E=m g h.
In physics, the law of conservation of energy governs that energy can neither be created nor destroyed, Potential Energy (P.E.) at the start of a fall must be equal to the Kinetic Energy (K.E.) at the end of the fall, which is expressed as the product of one half mass (½m) and impact velocity squared (v2). Therefore P.E. = K.E. = ½ mv2, Solving for linear impact velocity gives Equation 5:
Equation 5 – Linear impact velocity:
5.12. Angular vertical impact velocity
In real-world scenarios a falling rider will not follow a purely linear path [19], especially when coupled to a rigid body such as a motorcycle, hence angular velocity will also be generated (Figure 6).
Figure 6 – Falling Motorcyclist
If a motorcyclist falls from a vertical to a horizontal position, we can assume that Potential Energy (P.E.) is converted to rotation: 1/2m g h = ½ I ω2 where is the Moment of inertia, defined as the ratio of the angular momentum (L) of a system to its angular velocity (ω) around an axis: I=L/w which may also be expressed in terms of its mass (m) and its distance (r) from the pivot point as: I=mr2. Since r = h, the equation can be rewritten: mgh=1/2mh2w2. Instantaneous angular velocity at impact can be expressed in terms of linear components: ν = ω h, thus mgh=1/2mv2 which yields Equation 6:
Equation 6 – Instantaneous velocity due to angular rotation upon impact:
Hence total impact velocity is the sum of its linear and angular components.
Therefore, the sum of impact velocity due to linear and angular components is greater than impact velocity due to linear components only and is expressed as:
Equation 7 – Impact velocity:
5.13. Travel impact velocity
As previously stated, air resistance during a short fall is considered negligible, therefore angle and velocity at ejection (α, νejection) is considered equal to the angle and velocity at impact. Velocity due to ejection can be expressed in terms of its vertical and horizontal components . Assuming that ejection angle is measured with reference to the horizontal axis, then:
Equation 8a – Vertical ejection velocity: , and Equation 8b – Horizontal ejection velocity:
5.14. Impact velocity vector
The impact velocity vector has both vertical and horizontal components. The total vertical velocity is the sum of the linear and angular velocity components, plus the vertical components of velocity due to ejection. The total horizontal velocity will equal the horizontal component of velocity due to ejection. The magnitude of the impact velocity vector will be the square root of the sum of its vertical and horizontal components, hence:
Equation 9 – Impact velocity vector:
and the effective angle of the impact velocity vector relative to the vertical axis is determined as:
Equation 10: Effective impact angle:
5.15. Impact acceleration
In a motorcycle accident reconstruction, impact acceleration is determined as a function of rate of change of impact velocity over time (t): , where the duration of the impact will be directly affected by the stopping distance of the impacted material. Roadside materials, such as grass or dirt inherently have larger stopping distances than typical roadway materials, such as asphalt or concrete. Hence, the impact accelerations experienced by a rider landing on a grassy area will be considerably less than if they impacted the roadway.
Motorcycle Accident Reconstruction Case Study
A cruiser motorcycle was traveling along a divided highway, approaching an intersection, when a slow-moving automobile made an abrupt unanticipated lane change immediately in front of the motorcycle. The rider applied the brakes, locking up the rear wheel, causing the motorcycle to skid. The motorcyclist swerved in an attempt to avoid contact with the automobile. The left motorcycle footplate struck the rear corner of the automobile at an impact speed of 7.2 m/s (16 mph), causing the motorcycle to rotate violently about its long axis until the tires gained traction and the rider was thrown from the motorcycle. The final resting position of the rider was 4.6 m (15 ft.) past the final resting position of the motorcycle, to which the rider slid approximately 3.7 m (12 ft.) after being vaulted approximately 6.1 m (20 ft.) from the motorcycle (Figure 7).
Figure 7 – Case study: Automobile Avoidance Collision
The rider center of mass was calculated based on anthropometric derivation from known standing height of 1.7 m (5’8”) and manufacturer’s seat height specification of 0.7 m (27.5”), from which a suspension compression factor of 30 mm (1.2”) was subtracted. Head center of mass was calculated to be 1.45 m (57”). Given a minimal back angle correction factor, based on rider position on a cruiser motorcycle, the corrected HCOM was 1.4 m (55”).
Based on the distance that the rider was thrown and given an ejection velocity of 16 mph an ejection angle of 42 degrees was computed in this motorcycle accident reconstruction. Hence, it was determined that the rider gained an additional height of 1.2 m (39”) due to ejection, which is added to the rider head center of mass height of 1.4 m (51”), for a total fall height of 2.6 m (8’6”). Using equation 7, an impact velocity of 8.0 m/s (18 mph) was calculated for the rotating fall. Since impact velocity and angle is assumed identical to ejected velocity and angle, travel velocity expressed in terms of its vertical and horizontal components, are 4.8 m/s (10.7 mph) and 5.3 m/s (11.9 mph), respectively. Therefore, the total impact velocity vertical and horizontal components are (8.0 + 4.8) = 12.8 m/s (28.7 mph) and 5.3 m/s (11.9 mph), respectively, with an effective impact angle of 22 degrees relative to the vertical axis.
The helmeted motorcyclist impacted an asphalt roadway, head first. Given the inherently very short stopping distance of such materials, the duration over which the impact velocity was experienced was very short, resulting in high impact accelerations, which produced life-threatening traumatic brain injuries.
The results computed by our motorcycle accident reconstruction model were validated by and corroborated based upon physical evidence from the accident scene as well as the physical evidence of the injuries sustained by the rider.
Conclusions
The motorcycle accident reconstruction model presented herein has been successfully applied to a typical case study involving a single motorcycle collision. Measures of rider anthropometry were incorporated into the model. In the presented motorcycle accident reconstruction case study, the rider’s stature was smaller than that of an average male and seat height was lower than most stock motorcycles. Had average male stature and average motorcycle seat height been utilized, such assumptions would have over-estimated total fall height, thereby producing a calculated vertical impact velocity greater than was actually realized. In certain circumstances, specifically where ejection angle approaches 45 degrees, a simplified model without correction for rider anthropometry and rider posture might produce results that are in disagreement with physical evidence from the accident scene. However, this improved model is not without limitations. Specifically, if a rider were leaning the motorcycle considerably at the time of loss of control, such as when cornering, the initial vertical component (yo) would be reduced. This lean angle could be estimated given the radius of the corner and if the initial speed of the motorcycle can be computed. Overall, the validation of our new motorcycle accident reconstruction model is demonstrated in its application to the motorcycle accident reconstruction case study, which is in agreement with physical evidence from the accident scene.
8. References
[1] RnR Market Research (2014) Market Research Reports Press Release: Global motorcycles market demand to rise 7.2% annually to 2016. Published July 31.
[2] National Safety Council (2013) Injury Facts – 2013 Edition. Itasca, IL..
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* 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.
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Biomechanics (1899) is derived from the Ancient Greek bios “life” and mēchanikē “mechanics”, to refer to the study of the mechanical principles of living organisms, particularly their movement and structure. The earliest known reference to the study of biomechanics dates back to Aristotle (384– 322 BC), who published ‘DeMotu Animalium’ (On the Motion of Animals), in which he presented the mechanical concept ‘Ground Action Force’ as a starting point to deliberate where movement comes from.
The science of biomechanics has come a long way since the days of Aristotle. Contemporary biomechanics involves the application of Newtonian mechanics to determine physical capabilities and limitations of the human body. Trauma biomechanics examines whether mechanical forces acting on and within the human body may be sufficient to cause injury. The science of biomechanics is highly accepted by the courts for the purpose of explaining the mechanical causation of injuries.
Biomechanists posses advanced knowledge of human anatomy, mathematics and physics. We use this knowledge to study failure thresholds of human tissue, bone, ligaments, blood vessels, etc. When applying this knowledge to litigation, a biomechanist will perform a reconstruction to determine the forces acting on the plaintiff during the claimed injury-causing event and relate those forces to thresholds of injury. Biomechanists and Medical Doctors serve complementary roles in the medico-legal system. However a biomechanist is uniquely qualified, based on education, training and experience, to determine injury causation.
The methods that I use in my biomechanical evaluations are similar to methods that have been employed by other researchers and are generally accepted by experts in my field. Such methods have been validated and published in peer-reviewed scientific journals.
Expert in Injury Biomechanics
Dr. John Lloyd has served as a biomechanics expert for both defense and plaintiff’s counsel on hundreds of cases throughout the United States involving automobile collisions, motorcycle accidents, trucking crash as well as slips trips and falls. Dr. Lloyd is available to travel to investigate the causes of such cases. Based on his doctorate in ergonomics with a specialization in biomechanics, Dr. Lloyd can assess whether the claimed injuries meet or exceed known biomechanical thresholds of injury.
Biomechanics Laboratory
I utilize a state-of-the-science biomechanics laboratory in my evaluations, as depicted in the following figure. This biomechanics laboratory includes various certified biofidelic mannequins, dedicated test apparatus, data acquisition hardware, software, calibrated sensor instrumentation, professional photography, and high-speed and videography equipment.
Much of my research and work focusses on biomechanical evaluation of helmets, in particular motorcycle and sports helmets, including football and ski helmets.
For helmet testing, we have a standard NOCSAE (National Operating Committee for Standards in Athletic Equipment) head drop system
However, the standard NOCSAE system only measures forces associated with linear acceleration, which are attributed with focal head injuries, such as skull fractures. This system has a rigid neck and therefore cannot measure rotational or angular accelerations, which are associated with traumatic brain injuries, such as concussion and subdural hematomas. We have a modified helmet drop test system, developed in collaboration with the University of Maine, Advanced Manufacturing Center, validation of which has been published in a peer-reviewed journal.
Additionally, the biomechanics laboratory is equipped with the following resources:
Monorail head drop assembly
Twin wire guided drop system (NOCSAE)
Weighted pendulum impactor
Linear bearing table
Height-adjustable, eletromagenetically-controlled freefall drop platform
20,000N impact force plate
880lb ceiling mounted lift system
Certified biofidelic adult headforms
CRABI12 biofidelic infant mannequin
Hybrid III 3-yr old biofidelic mannequin (KSS)
National Instruments 32 channel USB-6343 X-series data acquisition system
LabView 2009 data acquisition software.
Calibrated sensors, including Kistler and PCB Piezotronics tri-axial accelerometers, MEMS triple axis digital gyroscopes, and PCB Piezotronics uni-axial and tri-axial load cells.
Selection of flooring materials, including carpeting, wood and laminates as well as concrete and wood sub-flooring surrogates
Professional still photography equipment
Normal speed and high speed (up to 1kHz) videography equipment
Photography flash and ‘hot’ lighting
Please call Dr. Lloyd at 813-624-8986 or email DrJohnLloyd@Tampabay.RR.com to discuss how he can be of assistance with your case.
Football helmet expert, Dr. John Lloyd, had the privilege to present his research on football helmets as part of the Keynote address at the National Instrument conference in Austin, TX this week. The audience of 5,000+ attendees learned about Dr. Lloyd’s research into biomechanics of the brain.
It has been said that helmets cannot prevent concussions. I disagree.
As a biomechanist I have dedicated my career to studying the biomechanics of brain injuries. There are two key mechanical forces that give rise to head and brain injuries (1) linear forces, which are responsible for visible injuries, including bruising and skull fractures, and (2) rotational forces, which cause invisible injuries, such as concussion and brain injury.
Since helmets are currently designed to pass testing standards that focus on linear forces only, it is no surprise that helmets have limited benefit in preventing concussions. Through advances in medicine we have learned that concussions can potentially have life-long neurological consequences, including memory impairement and personality changes / behavioral effects.
Over the past years I have developed and validated a testing method to evaluate helmets in terms of their ability to protect against both linear and rotational forces. Using this apparatus I characterized football helmets, results of which have been submitted to Science for publication.
Based on lessons learned from my biomechanical evaluation of various sports helmets, I have devised a matrix of shear-thickening non-Newtonian materials. A prototype helmet was constructed using this matrix liner, results of which show that rotational forces that cause concussion and other brain injuries are reduced by up to 50% compared to a leading football helmet, while also reducing linear forces.
helmet prototype reduces concussion risk
It is my goal and my passion to work with leading helmet companies to make this technology available to players and sports participants of all aged to enhance their protection against brain trauma. I am looking to collaborate with one manufacturer in each sport to offer an exclusive license patent-pending technology.
In December a movie titled “Concussion”, staring Will Smith will be released in theaters, chronicling the work and bravery of Dr. Bennett Omalu, who first discovered Chronic Traumatic Encephalopathy (CTE) as the consequence of repeated blows to the brain in football and attempts by the National Football League (NFL) to deny any causal link.
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 
