Biomechanics expert Dr. John Lloyd has served attorneys nationwide for 25+ years in biomechanics, human factors, helmet testing and motorcycle accident expert
Dr. Lloyd is a distinguished authority in motorcycle accident reconstruction and human factors analysis, with decades of experience. His understanding of the unique dynamics involved in motorcycle crashes sets him apart as a true specialist in the field.
Dr. Lloyd spent his career as a senior researcher at the VA Hospital in Tampa, FL, serving as Director of the Biomechanics Research Laboratory and Director of the Traumatic Brain Injury Research Laboratory. In addition he held a courtesy faculty appointment as Assistant Professor in the University of South Florida College of Engineering from 2002-2022, and is currently the Research Director of BRAINS, Inc.
To date, Dr. Lloyd’s work has been published in six book chapters and 33 peer-reviewed journals, as well as presented at more than 100 national and international conferences (see curriculum vitae).
Comprehensive Approach
Dr. Lloyd goes beyond the obvious and delves deep into the technical intricacies of each case. As a multi-disciplinary expert he combines, accident reconstruction, biomechanics and human factors to provide a holistic view of the accident, ensuring no detail goes unanalyzed.
Accurate Motorcycle Crash Reconstructions
Using state-of-the-science reconstruction tools and real world data, Dr. Lloyd meticulously creates 3D accident reconstructions with unparalleled accuracy. This empowers him to provide precise insights into the sequence of events leading up to the incident.
Human Factors Insight
Understanding the role of human behavior is crucial in accident analysis. Dr. Lloyd’s human factors expertise allows him to investigate the cognitive factors affecting both motorcycle riders and automobile drivers, offering invaluable insights into decision-making processes.
Courtroom Excellence
Dr. Lloyd’s reputation as a credible and authoritative expert makes him an invaluable asset in the courtroom. He excels at conveying complex technical information to the jury in an accessible manner, helping you present a compelling case, backed by robust scientific analysis.
To date, Dr. Lloyd has provided expert witness Deposition and Trial Testimony in more than 160 civil and criminal cases. His expertise in motorcycle crashes, motorcycle riding and operation, helmet protection, biomechanics and human factors has been recognized by courts across the United States and Internationally. The analysis methods that Dr. Lloyd utilizes are published in peer-reviewed scientific journals.
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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:
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..
[4] NHTSA’s National Center for Statistics and Analysis (2007) Motorcycles Traffic Safety Fact Sheet (DOT-HS-810-990), 1200 New Jersey Avenue SE, Washington, DC 20590: National Highway Traffic Safety Administration.
[5] Hurt HH, Ouellet JV, Thom DR (1981) Motorcycle Accident Cause Factors and Identification of Countermeasures. Volume 1: Technical Report. University of Southern California Traffic Safety Center, Los Angeles, CA..
[6] ACEM (2000) MAIDS (Motorcycle Accidents In Depth Study): In-depth investigations of accidents involving powered two wheelers – Final Report. European Association of Motorcycle Manufacturers, Brussels..
[7] U.S. Government Accountability Office (2012) Motorcycle Safety: Increasing Federal Funding Flexibility and Identifying Research Priorities Would Help Support States’ Safety Efforts. Report number GAO-13-42.
[9] Medwell C, McCarthy J, Shanahan M (1997) Motorcycle Slide to Stop Tests. SAE Technical Paper 970963., SP-1237 Accident Reconstruction and Animation VII, Warrendale, PA
[10] Southwestern Association of Technical Accident Investigators (1984) Motorcycle Drag Factor Tests. Phoenix, AZ.
[11] Day TD, Smith JR (1984) Friction Factor for Motorcycles Sliding on Various Surfaces. SAE paper 840250. Society of Automotive Engineers, Warrendale, PA.
[12] Iowa State Patrol (1985) Motorcycle Test Skidding on its Side, Traffic Investigation Spring Seminar. Johnston, IA..
[13] Royal Canadian Mounted Police (1984) Motorcycle Testing. Coquitlam, BC, Canada.
[14] Hague DJ (2001) Calculation of Impact Speed from Pedestrian Slide Distance. Proceedings of The Institute of Traffic Accident Investigators International Conference
[15] Searle JA, Searle A (1983) The Trajectories of Pedestrians, Motorcycles, Motorcyclists, etc., Following a Road Accident. SAE paper 831622.. Society of Automotive Engineers, Warrendale, PA.
[16] Thede P, Parks L (2010) Race Tech’s Motorcycle Suspension Bible. Motorbooks International publisher, UK. Cd s.
[17] Pheasant, S. (1998) Bodyspace. Taylor and Francis, London.
[18] Gordon CC, Churchill T, Caluser CE, Brandtmiller CB, McConville JT et al. (1989).1988 Anthropometric Survey of US Army Personnel. US Army Technical Report TR-89/044. Natick, MA.
[19] Barnett, RL (1995) The drunk, the child and the soldier – my how they fall. Triodyne Inc. Safety Bulletin. ISSN 1081-4140. Vol 2 (2).
Motorcycle collision analysis is a highly specialized discipline in which Dr. Lloyd is eminently qualified as a motorcycle accident expert. In addition to holding a PhD in Ergonomics (Human Factors), with a specialization in Biomechanics, John has more that 20 years and 200,000 miles of experience riding motorcycles. Dr. Lloyd has completed numerous advanced programs, including Motorcycle Safety Foundation (MSF), Experienced Rider Course and Total Rider Tech Advanced training.
Motorcycle Helmets and Brain Injury
To consider whether a motorcycle helmet might reduce the risk of brain trauma in a motorcycle accident it is first important to understand the two primary mechanisms associated with traumatic brain injury – impact loading and impulse loading.
Impact loading involves a direct blow transmitted primarily through the center of mass of the head, resulting in extracranial focal injuries, such as contusions, lacerations and external hematomas, as well as skull fractures. Shock waves from blunt force trauma may also cause underlying focal brain injuries, such as cerebral contusions, subarachnoid hematomas and intracerebral hemorrhages. Whereas, impulse or inertial loading caused by sudden movement of the brain relative to the skull, produces cerebral concussion. Inertial loading at the surface of the brain can cause subdural hemorrhage due to bridging vein rupture, whereas if affecting the neural structures deeper within the brain can produce diffuse axonal injury (DAI).
Holbourn was the first to cite angular / rotational acceleration as an important mechanism in brain injury. Gennarelli, Thibault, and colleagues, in a series of studies using live primates and physical models investigated the role of rotational acceleration in brain injury. They concluded that angular acceleration contributes more than linear acceleration to brain injuries, including concussion, axonal injury, and subdural hematoma.
Motorcycle Helmet Testing
Traditional testing of motorcycle helmets focuses 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. According to the Federal Motor Vehicle Safety Standard (FMVSS) 218, commonly known as the DOT helmet standard, the test involves dropping a motorcycle helmet onto a flat steel and hemispherical anvil at an impact velocity of 6.0 m/s (13.4mph). In general, if peak linear acceleration is less than 400g, the helmet is considered acceptable. Current motorcycle helmet testing standards do not incorporate measures of angular acceleration and therefore do not address whether any helmets can provide adequate protection against catastrophic brain injuries, such as concussion, axonal injury and subdural hematoma.
In 1995, the European Commission Directorate General for Energy and Transport initiated a Cooperative Scientific and Technical Research (COST) program to investigate Motorcycle Safety Helmets. Several agencies from Finland, the United Kingdom, France and Germany participated in this study, which compiled and analyzed data from 4,700 motorcycle fatalities in Europe, each year. The COST report documents that 75% of all fatal motorcycle accidents involve head injury. Linear forces were present in only 31% of fatal head injuries, while rotational forces were found to be the primary cause in over 60% of cases. Within the scope of this study experiments were performed using drop tests with accelerometers to measure linear and rotational accelerations of the brain and skull mass associated with different types of impacts. These tests confirmed rotational acceleration to be a primary cause of brain injury in helmeted motorcycle accidents.
Rotational forces acting on the brain are the underlying cause of traumatic brain injuries.
Motorcycle helmets, including those certified under DOT and SNELL standards are designed to mitigate forces associated with linear acceleration.
Motorcycle helmets are not currently certified under either DOT or SNELL standard against their ability to protect against the angular / rotational forces.
Epidemiologic evidence from the COST-327 report indicates that motorcycle helmets do not provide adequate protection against closed head and brain injuries
Human Factors of Motorcycle Accidents
Human factors in vehicle collisions include all factors related to drivers and other road users that may contribute to a collision. Examples include driver behavior, visual and auditory acuity, decision-making ability, and reaction speed. A 1985 report based on British and American crash data found driver error, intoxication and other human factors contribute wholly or partly to about 93% of crashes.
Motorcycle Inspection
Motorcycle accident analysis often requires involves a teardown and careful inspection of the machine to investigate for possible contributing factors. Our engineers have a combined 70 years experience with motorcycle mechanics.
A thorough evaluation includes inspection of tires, brakes, suspension setup, electrical components as well as any aftermarket parts.
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.
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.
John Lloyd of BRAINS, Inc. announced today that football head injuries and concussions can be reduced up to 50 percent with their new helmet safety breakthrough.
San Antonio, FL – Dr.John Lloyd PhD 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 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, Inc. 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 up to 50% compared to similar testing with a leading football helmet,” said Dr. John Lloyd, Research Director.
“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.
About BRAINS, Inc.
BRAINS, Inc. located in San Antonio, Florida, is a research and development company focused on the biomechanics of brain injuries. The company was founded in 2011 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. BRAINS, Inc. 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 visit : http://drbiomechanics.com/sports-helmet-football-helmets/new-helmet-technology/ or call 813-624-8986.
Dr. Lloyd’s research article “Brain Injury in Sports”, co-authored with Dr. Frank Conidi has been published in the Journal of Neurosurgery.
Please email me at DrJohnLloyd@Tampabay.RR.com if you would like to receive a full copy of the published article.
Abstract
BACKGROUND
Helmets are used for sports, military, and transportation to protect against impact forces and associated injuries. The common belief among end users is that the helmet protects the whole head, including the brain. However, current consensus among biomechanists and sports neurologists indicates that helmets do not provide significant protection against concussion and brain injuries. In this paper the authors present existing scientific evidence on the mechanisms underlying traumatic head and brain injuries, along with a biomechanical evaluation of 21 current and retired football helmets.
METHODS
The National Operating Committee on Standards for Athletic Equipment (NOCSAE) standard test apparatus was modified and validated for impact testing of protective headwear to include the measurement of both linear and angular kinematics. From a drop height of 2.0 m onto a flat steel anvil, each football helmet was impacted 5 times in the occipital area.
RESULTS
Skull fracture risk was determined for each of the current varsity football helmets by calculating the percentage reduction in linear acceleration relative to a 140-g skull fracture threshold. Risk of subdural hematoma was determined by calculating the percentage reduction in angular acceleration relative to the bridging vein failure threshold, computed as a function of impact duration. Ranking the helmets according to their performance under these criteria, the authors determined that the Schutt Vengeance performed the best overall
CONCLUSIONS
The study findings demonstrated that not all football helmets provide equal or adequate protection against either focal head injuries or traumatic brain injuries. In fact, some of the most popular helmets on the field ranked among the worst. While protection is improving, none of the current or retired varsity football helmets can provide absolute protection against brain injuries, including concussions and subdural hematomas. To maximize protection against head and brain injuries for football players of all ages, the authors propose thresholds for all sports helmets based on a peak linear acceleration no greater than 90 g and a peak angular acceleration not exceeding 1700 rad/sec2.
Please call Dr. Lloyd at 813-624-8986 or email DrJohnLloyd@Tampabay.RR.com if you would like to receive a full copy of the published article “Brain Injury in Sports”
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.
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.
Opportunity to Protect Professional and Youth Sports Players from Traumatic Brain Injuries
Sport concussion researchers teamed up with football players at a Florida high school. Ten players were equipped with Riddell Revolution Speed helmets, with the embedded Simbex HITS encoders, which were worn throughout the 2011/2 football season. The HITS system recorded the severity and location of all head impacts during both football practice sessions and games.
To measure the physiological effects of acute and cumulative head impacts, players agreed to wear a wireless EEG system, which was housed on the back of the shoulder pads. In addition, heart rate variability, respiration rate as well as linear and angular motion was recorded using a Tricorder developed by ReThink Medical.
During the 2011/2- football season, several concussive level impacts were recorded. Two players were removed from the field due to suspected sport concussion / mTBI, one of whom was wearing the complete data acquisition system, including HITS encoders, Nicolet EEG and ReThink Tricorder at the time of impact and for approximately 30 minutes post-impact. For the first time we have the opportunity to investigate physiological responses and brain activity changes in response to a concussive level head impact.
Analysis of one player’s self-reported concussive impact clearly shows decreased Gamma band activity and increased Theta band activity in the frontal cortex of the brain immediately following significant head impact. This suggests that the player had reduced cognitive performance and was perhaps in a ‘drowsy’ state for about 10 minutes following impact. During this time, the player may have been dazed and confused and certainly less effective on the field. But more importantly, his ability to protect himself from a second, potentially harmful impact was greatly compromised.
The findings of our study clearly indicate compromised brain activity as a result of head impact, which appears to be correlated with the magnitude of the impact.
Normalized Power Trend Analysis. Normalized Theta (Left) and Gamma (Right) Power (log of % power within band) of a football player, who experienced a concussion following a moderately forceful head impact (Red line), show phasic modulations in power throughout the practice. Fluctuations in power rarely exceed 25% of the total average power for the recording session in Theta and Gamma frequencies. Yet, immediately following a violent hit (Red line), gamma power begins to decline rapidly and exceeds an arbitrary criterion of ±50% change from average power (peaking at 90 min.). Indeed gamma power remained within 20% of the mean for most of the duration of practice, exceeding this degree of change for over 10 minutes after the impact and two other brief episodes (around 20 min. and 50 min. for less than five minutes; Note, the first and last five minutes were ignored due to the temporal filtering artifact at both edges). Whereas, a peak in theta power coincided with the greatest change in gamma power, the degree of change from the mean normalized power never exceeded 10%. This preliminary data suggests that our algorithms provide (1) the sensitivity to detect significant change in brain activity following a concussive event, and (2) specificity in detecting which frequency band (i.e., gamma) provides the most meaningful brain signal for detecting concussion / brain trauma
Our future goals for the upcoming football season include a new micro-EEG recorder, which is in development, that will allow unobtrusive measurement of several players simultaneously during both football practice and games.
Ultimately, it is our hope that this technology will be widely available to both professional and youth teams so that medical staff can monitor the brain health of players in real-time so that injured participants can be objectively identified, effectively protected and successfully treated.
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.