Expert witness Dr. John Lloyd has served attorneys nationwide for 25+ years in biomechanics, human factors, helmet testing and motorcycle accident expert
Crash data retrieval (EDR) is a tool used by accident reconstructionists to download pre and post collision information from vehicles to determine driver behavior, impact severity and the cause of the crash.
Much like an airplane, almost all vehicles in the United States are equipped with a ‘black box’. In the event of an impact critical information about the crash as well as pre-crash information is recorded and stored in this black box, also known as an event data recorder (EDR). The process of obtaining this data is known as crash data retrieval.
It is a common misunderstanding that EDR data is only temporarily saved in the vehicle’s ‘black box’ and is erased after a short period of time. In the event of an airbag deployment, EDR data is almost always recorded and permanently stored. This data is never erased. Whereas, in a non-deployment event (i.e. airbags did not go off), if the event meets the manufacturer’s propriety algorithms to trigger recording, that data is saved, but can be overwritten if the vehicle is involved in subsequent collisions.
This information of vital importance when investigating a crash includes:
Speed
Change in Velocity (Delta-v)
Acceleration / Deceleration
Braking
Pre-crash Steering Inputs/ Angle
Seatbelt Usage and airbag deployment
Specialized equipment is required to access the EDR.
Unfortunately, law enforcement investigating a crash typically only retrieve the information from the Crash Data Recorder in the event of a traffic homicide. Hence, it is up to the accident reconstruction experts to download this information from the vehicles.
EDR data can help an accident reconstructionist to determine:
pre-crash driver behavior
crash severity
who was at fault
Combined with his biomechanical expertise, Dr. Lloyd can use the information from the EDR to determine:
occupant kinematics (motion in response to crash)
whether injuries are consistent with crash severity
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.
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..
[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.
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On the admissibility of biomechanics testimony the Reference Manual on Scientific Evidence, written jointly by the National Academies of Sciences, Engineering and Medicine and the Federal Judicial Center, states: “Specifically, one cross-disciplinary domain deals with the study of injury mechanics, which spans the interface between mechanics and biology. The traditional role of the physician is the diagnosis (identification) of injuries and their treatment, not a detailed assessment of the physical forces and motions that created injuries during a specific event. The field of biomechanics (also called biomechanical engineering) involves the application of mechanical principles to biological systems, and is well suited to answering questions pertaining to injury mechanics.” In the case Garner v Baird [910 N.Y.S.2d 762, 762 (N.Y. Civ. Ct. 2010)] defined biomechanics as “the application of physics and mechanical engineering to the human body.”
In a ruling of the 1st District Court of Appeals of Florida on July 19, 2012 [98 So.3d 115, Florida First District Court of Appeals, 2012] Judge Healey concluded, “that biomechanics expert, Dr. John Lloyd is qualified to offer opinions as to causation because the mechanism of injury fell within the field of biomechanics”. Moreover, in the case of Taylor v Culver Florida First District Court of Appeals, 2015 the appeals court ruling, which directly references Council states “the proffered testimony of the Appellant’s biomechanics expert was relevant to the disputed issues concerning velocity and direction of forces involved in the accident”. In the case Maines v Fox [190 So.3d 1135, Florida First District Court of Appeals, 2016], the ruling states: “Biomechanical opinions as to the general causation of a type of injury are admissible in a personal injury case.”
* 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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[44] Rowson S, Duma SM (2013) Brain injury prediction: assessing the combined probability of concussion using linear and rotational head acceleration. Ann Biomed Eng. 41(5): 873-82. PMID: 23299827
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[47] Mellor AN, St Clair VJM, Chinn BP (2007) Motorcyclists’ helmets and visors – test methods and new technologies. TRL Limited, Wokingham, Berkshire, UK. Project # S0232/VF.
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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”
Distracted driving is any activity that diverts attention from driving, including talking or texting on your phone, smoking, eating and drinking. I have even seen people putting on makeup and shaving their face while driving!
Texting is the most alarming driver distraction. Sending or reading a text takes your eyes off the road for 5 seconds or more. At 40 miles per hour a car can travel the length of a football field in 5 seconds; at highway speeds, a vehicle can travel the length of almost two football fields in the same time.
Texting while driving and other cell-phone reading and writing activities are high-risk activities associated with motor vehicle collisions and mortality. In 2015 alone 3,450 people were killed on US roads. 391,000 were injured in motor vehicle crashes involving distracted drivers.
There are three types of driver distractions – manual, visual and cognitive.
Manual distractions cause you to take your hands off the steering wheel
Visual distractions cause you to take your eyes off the road
Cognitive distractions interfere with your mental focus on driving
Texting involves all three of the above distractions and is therefore considered a high risk activity while driving.
Many distractions exist while driving. Use of cellular phones while driving is a top distraction. Most new vehicles are equipped with hands-free technology, leading drivers to believe that talking via bluetooth is safe. However, this is not the case. In fact, human factors research shows that our brains remain distracted from the task of driving for several minutes even after hanging up a call. According to the National Safety Council, the use of cell phones while driving causes approximately 26% of motor vehicle crashes.
Human factors research shows that cell phone usage while driving is associated with increased reaction time, decreased situational awareness, more impulsive responses, and shrinks a drivers search area. Further, drivers exhibit less attention to objects even when looking directly at them. Moreover, “Impairments associated with using a cell phone while driving can be as profound as those associated with driving while drunk”. Hands free (Bluetooth) cell phone use is no less distracting than hand-held. What makes cell phone usage particularly dangerous is that drivers do not appreciate that their abilities are impaired due to distraction. Drivers distracted due to the use of cellular phones have little or no input to allow them to appreciate the things they did not see, or the things they looked at and did not process adequately.
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.
Specialized equipment is required to access the EDR.
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 
