Helmet expert Dr. John Lloyd has served attorneys nationwide for 25+ years in biomechanics, human factors, helmet testing and motorcycle accident expert
The common belief among riders is that a motorcycle helmet protects the whole head, including the brain. However testing standards in Europe (ECE 22.05) and the US (DOT & Snell), which involve dropping helmeted headforms from heights of 2-3 meters onto a steel plate, only evaluate a motorcycle helmet in terms of its ability to protect against blunt force trauma, such as skull fractures and penetrating head injuries. The mechanism underlying diffuse brain injuries, such as concussions and brain hemorrhages is distinctly different, but is not assessed by current motorcycle helmet testing standards.
Imagine a bowl of jelly, where the bowl represents the skull and the jelly represents the brain. The bowl (skull) serves to protect the jelly (brain) from impact by dispersing forces over a larger surface area. If the bowl were impacted such that the force passes through the center of the jelly, the jelly moves very little. This is called linear force. Whereas, if you rotate the bowl of jelly between your hands you will see that the jelly moves quite a lot, especially towards its center. This is called a rotational force.
In reality, most motorcycle helmet impacts will produce both linear and rotational forces. In the case of head and brain injury, linear forces are responsible for injuries such as bruises and fractures. Whereas rotational forces cause the nerves and blood vessels in the brain to stretch and tear, leading to concussions, injury to the nerve fibers (axonal trauma) and brain bleeding (hematomas).
The human head is designed to protect the brain against typical impacts associated with daily living, such as normal bumps and falls. The skull can be thought of as a helmet to the brain by resisting penetrating injury to the brain. While the scalp glides over the skull to decrease rotational forces, thereby reducing the risk and severity of diffuse brain injuries. However, the forces associated with motorcycle collisions far exceed that which the human skull and scalp was intended to protect. Hence in motorcycling the use of a helmet to reduce the risk of such injuries is typically mandated.
Helmets are designed with 3 principal components – the outer shell, the inner liner and a comfort layer. The shell is typically made of polycarbonate plastics or fiberglass and serves two purposes; to minimize the likelihood that a sharp object might penetrate the head, and to dissipate the impact over a larger surface area. The inner liner is made from EPS foam (polystyrene) and serves to absorb the impact forces. The comfort layer does nothing more than provide comfort between the head and the polystyrene liner. Unfortunately, the polystyrene liner has limited effectiveness at reducing the rotational forces – those responsible for diffuse brain injuries – below safe levels.
A cooperative study was undertaken in Europe in the late 1990s to examine motorcycle accidents and their causes. Based on data from 4,700 helmeted motorcyclist deaths, the study found head injuries accounted for three-quarters of all fatalities. More than 60 percent of which were brain injuries caused by rotational forces, while only 30 percent of fatal head injuries were due to linear forces. This extensive study proves that motorcycle helmets are inadequate in providing necessary protection against diffuse brain injuries.
One might propose that protection against diffuse brain injury ought to deserve a higher priority. After all, the skull will likely heal from trauma, but the brain may not.
The challenge with protective headgear, including motorcycle, military and sports helmets is that, due to the characteristics of the liner materials, the head is directly coupled to the helmet. That is, the head and helmet are effectively joined and move as one. Therefore upon impact, any rotational forces generated on the helmet are transmitted directly to the brain. In fact, due to the size of helmets rotational forces can actually be amplified. The solution lies in de-coupling the head from the helmet, much the way that the scalp is de-coupled from the skull, so that the helmet can have some degree of rotation independent of the head. In this way, the rotational forces are dampened before they are transmitted to the brain, thereby lessening the risk and severity of brain injury.
BRAINS, Inc., of which Dr. Lloyd is the Research Director, is developing a new generation of motorcycle helmets, utilizing a patented composite of shear-thickening non-Newtonian materials. Due to their nature, these advanced materials respond differently to linear and rotational forces, thereby allowing the helmet some independent rotational motion, effectively de-coupling the helmet from the head. This technology was demonstrated at NI Week (http://youtu.be/T591x950oRI) and shows great promise for protection against both blunt force trauma and traumatic brain injuries.
Given the choice of a helmet that protected against skull fracture and one which also provides protection against brain injury, which would you choose?
Dr. John Lloyd holds a PhD in Ergonomics from Loughborough University and is a Brain Injury Specialist. He is an expert in the field of brain injury biomechanics.
As a motorcycle enthusiast, John has clocked more than 250,000 miles and completed numerous training programs. Dr. Lloyd has served as a biomechanics expert on a variety of motorcycle accident cases.
Two helmeted motorcyclist were traveling on a rural state road when a tractor-trailer driver failed to see the bikes and made a left turn in front of them to enter a truck stop. The rider in the right track had little time to respond and collided head first into the box trailer. He was pronounced deceased at the scene.
The helmeted motorcyclist was wearing a non-compliant or ‘novelty’ helmet, which did not meet DOT motorcycle helmet standards (FMVSS 218). Opposing counsel claimed that had the biker been wearing a DOT-certified motorcycle helmet he may have survived the impact.
Motorcycle helmet expert, Dr. John Lloyd, was retained to evaluate and compare the protective performance of DOT-certified and novelty motorcycle helmets.
Based on a comprehensive motorcycle accident reconstruction it was determined that the impact speed of the rider was 45 to 50 miles per hour. Motorcycle helmet certification tests typically involve impact speeds of 13-17 miles per hour. Therefore a dedicated apparatus was constructed to generate higher impact speeds. Using a force-balanced twin pendulum apparatus, Dr. Lloyd was able to generate head impact speeds similar to those specific to the subject crash, yet preserve the standard DOT test methodology, thereby avoiding a Daubert challenge.
Eight DOT and non-DOT helmets were purchased for this study. Each was impacted once in the frontal region while fitted to an instrumented crash test dummy head. High speed data and video were acquired for each test.
Results demonstrate that, although the tested DOT-certified motorcycle helmets outperformed the tested novelty helmets, neither would provide adequate protection against head injuries, such as skull fractures, contusions and lacerations, or brain injuries, including hemorrhages or axonal injury in an impact of this magnitude.
Dr. Lloyd’s prior published motorcycle helmet studies demonstrate that while DOT-certified motorcycle helmets can reduce the risk of traumatic head injuries, typical helmets do not afford any protection against acute brain injury.
* 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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[11] Chang L, Chang G, Huang J, Huang S, Liu D et al. (2003) Finite Element Analysis of the effect of motorcycle helmet materials against impact velocity. Journal of the Chinese Institute of Engineers. 26: 835–843.
[16] Caccese V, Lloyd J, Ferguson J (2014) An Impact Test Apparatus for Protective Head Wear Testing Using a Hybrid III Head-Neck Assembly. Experimental Techniques. PMID: 28216804
[17] Lloyd J & Conidi F (2015) Brain Injury in Sports. Journal of Neurosurgery. 124(3):667-74 PMID: 26473777
[22] BSI (1985) BS 6658 – Specification for protective helmets for vehicle users. British Standards Institute.
[23] ECE (2002) 22.05 Protective Helmets and their Visors for Drivers and Passengers of Motorcycles and Mopeds.
[24] Australian/New Zealand Standard (2006) AS/NZS1698, Protective Helmets for Vehicle Users. Australian/New Zealand Standard.
[25] Ono K (1998) Human head impact tolerance. In Yoganandan (Ed). Frontiers in Head and Neck Trauma: Clinical and Biomechanical. IOS Press, Amsterdam.
[26] Nahum, A. M., Gatts, J. D., Gadd, C. W., Danforth, J (1993) Impact Tolerance of the Skull and Face. In Biomechanics of Impact Injury and Injury Tolerances of the Head-Neck Complex. Ed. Stanley H. Backaitis. Warrendale: Society of Automotive Engineers. 631-645.
[28] Gennarelli TA, Thibault LE, Adams JH, Graham DI, Thompson CJ, et al (1982) Diffuse axonal injury and traumatic coma in the primate. Ann Neurol. 12(6): 564-74. PMID: 7159060
[29] Gennarelli T. and Thibault L (1982) Biomechanics of Acute Subdural Hematoma, J Trauma. 22(8), 680-686. PMID: 7108984
[30] Lowenhielm P (1974) Dynamic properties of the parasagittal bridging veins. Z. Rechtsmed. 74 (1): 55-62. PMID: 4832079
[31] Lowenhielm P (1975) Strain Tolerance of the Vv. Cerebri sup. (Bridging Veins) Calculated from Head-on Collision Tests with Cadavers. Z. Rechtsmed. 75 (2): 131-144. PMID: 4217056
[32] Lowenhielm P (1978) Tolerance level for bridging vein disruption calculated with a mathematical model. J Bioengineering. 2 (6): 501-507. PMID: 753840
[33] Lee MC and Haut RC (1989) Insensitivity of tensile failure properties of human bridging veins to strain rate: implications in biomechanics of subdural hematoma. J Biomechanics. 22 (6-7): 537–542. PMID: 2808439
[34] Meaney DF (1991) Biomechanics of acute subdural hematoma in the subhuman primate and man. University of Pennsylvania. PhD dissertation.
[38] Gennarelli TA, Adams JH, Graham DI (1981) Acceleration induced head injury in the monkey I: The model, its mechanistic and physiological correlates. Acta Neuropathol. Suppl. 7:23-25. PMID: 6939241
[39] Thibault LE and Gennarelli TA (1985) Biomechanics of diffuse brain injuries. Stapp Car Crash Conference. Twenty-Ninth Proceedings, SAE Paper No. 856022, New York.
[40] Pellman EJ, Viano DC, Tucker AM, Casson IR, Waeckerle JF (2003) Concussion in professional football: reconstruction of game impacts and injuries. Neurosurgery. 53(4): 799-812. PMID: 14519212
[41] Rowson S, Brolinson G, Goforth M, Dietter D, Duma S (2009) Linear and angular head acceleration measurements in collegiate football. J Biomech Eng. 131(6). PMID: 19449970
[42] Rowson S, Goforth MW, Dietter D, Brolinson PG, Duma SM (2009) Correlating cumulative sub-concussive head impacts in football with player performance. Biomed Sci Instrum. 45:113-8. PMID: 19369749
[43] Rowson S, Duma SM, Beckwith JG, Chu JJ, Greenwald RM, et al (2012) Rotational head kinematics in football impacts: an injury risk function for concussion. Ann Biomed Eng. 40(1): 1-13. PMID: 22012081
[44] Rowson S, Duma SM (2013) Brain injury prediction: assessing the combined probability of concussion using linear and rotational head acceleration. Ann Biomed Eng. 41(5): 873-82. PMID: 23299827
[45] Adamson KS, Alexander P, Robinson EL, Johnson GM, Burkhead CI, et al. (2002) Seventeen motorcycle crash tests into vehicles and a barrier. SAE 2002-01-0551. Society of Automotive Engineers, Warrendale, PA.
[46] Severy DM, Brink HM, Blaisdell DM (1970) Motorcycle collision experiments. SAE Technical Paper 700897. Society of Automotive Engineers, Warrendale PA.
[47] Mellor AN, St Clair VJM, Chinn BP (2007) Motorcyclists’ helmets and visors – test methods and new technologies. TRL Limited, Wokingham, Berkshire, UK. Project # S0232/VF.
[48] Mertz HJ, Patrick LM (1971) Strength and Response of the Human Neck. Stapp Car Crash Conference; SAE Technical Paper 710855. Society of Automotive Engineers, Warrendale PA.
Football helmet research presented by John Lloyd, PhD – BRAINS, Inc. & Frank Conidi, MD – Florida Center for Headache and Sports Neurology at the 66th Annual Meeting of the American Academy of Neurology, 2014
Sports related concussion is the most widely publicized neurological disorder, with football accounting for the highest incidence across all sports. There is a silent epidemic of these invisible injuries across players of all ages from youth through professional, resulting in a 3-fold increase in ER visits among high school players from 2000-2010. It is estimated that a quarter of a million of these injuries each year have long-term consequences. A single moderate to severe brain injury can leave one at risk for early onset dementia, while repeated mild concussions may have the same effects (Giza).
The football helmet became mandatory in the 1930’s to provide protection against catastrophic head injuries. Over the past eighty years there have been significant modifications in football helmet design. Yet despite advances in technology there is still little evidence that helmets offer significant protection against concussion and traumatic brain injury (Giza). hile the widely utilized Simbex HITS system and the Virginia Tech STAR rating system attempt to measure helmet performance, neither offer a direct measure of concussion or brain injury risk.
Methods
According to Holbourn, risk of focal head injury, such as skull fracture and brain contusion, can be expressed in terms of linear acceleration, while the risk of concussion, axonal injury and diffuse brain injury is associated with angular/rotational acceleration. The standard NOCSAE / ASTM helmet tests only measure forces associated with linear acceleration and therefore fail to account for risk of brain injury.
A modified test apparatus, incorporating a Hybrid III crash test dummy head and neck, has been validated by Caccese and Lloyd. This method induces a rotational inertia on impact, thereby facilitating measurement of risk of focal head and diffuse brain injuries.
We purchased 60 football helmets, including three samples each of 20 different models. All helmets were dropped five times from a height of 2.0 meters onto a steel plate, generating an impact velocity of 13.9 mph.
Football Helmet Results
Our findings show that football helmets vary widely in terms of their performance to protect against focal head injury and concussion / diffuse brain injury. All tested football helmets, including the 1930s leatherhead meet the minimum performance criteria of 275G as set by the ASTM F717 standard. However, this standard does not account for duration of impact, which is a critical factor. Moreover, ASTM F717 does nothing to set helmets performance standards in terms of protection against concussion and brain injury.
By design, helmets reduce impact force by increasing the impact duration. But, as demonstrated by Depreitere (blue) and Lowenhielm (red) increasing impact duration actually lowers the brain injury threshold. Head injury risk was calculated with respect to the140g threshold for 7msec impact documented by Ono, wwhile brain injury risk was calculated based on Ommaya’s1700 rad/s^2 tolerance limit for moderate AIS2 brain injury, which concurs with Rowan and Duma’s top 25% of sub-concussive impacts.
Results are presented below, where % reduction of head injury risk is shown in red and % reduction of concussion/mTBI is presented in blue.
Bars above the x-axis indicate that the helmet performs better than the documented threshold, while those below the x-axis did not meet our performance thresholds.
It is noted, interestingly, that the 1930s Goldsmith leatherhead helmet actually outperformed several contemporary football helmets in terms of protecting against concussion and brain injury, including the Adams a2000 Pro, the Rawlings Quantum and the Riddell 360.
Based on our research, the top 3 varsity helmets are: Xenith X1, Schutt Air XP Pro, and Rawlings Quantum Plus, respectively. The top 10 helmets are presented below, based on their protection against traumatic focal head injuries and diffuse brain injuries:
Conclusions
None of the football helmets on the market today offer what most would consider adequate protection against concussions and traumatic brain injuries. A shift in thinking towards lighter high-tech materials for helmets, teaching proper hitting and tackling techniques, pre and in-season isometric and isokinetic cervical strengthening programs, and continued concussion awareness and education are the best means of protecting athletes of all levels from the consequences of concussion and traumatic brain injury.
Future Research
It is hypothesized that oblique impacts present even longer impact durations, which may explain why such seemingly innocuous impacts cause increased incidence of concussion.
Embracing the findings from testing of the 1930s Goldsmith leatherhead, a new generation of soft football helmets is proposed, utilizing today’s advanced non-Newtonian materials, which we anticipate might outperform contemporary helmets in terms of protection against both traumatic head and brain injuries.
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”
Helmets are intended to minimize blunt force trauma to the head, such as skull fracture, lacerations and contusions. Whereas risk of diffuse brain injuries, such as concussion, brain bleeding and axonal injuries are caused when brain tissue, nerves and blood vessels stretch and tear as the head moves suddenly but the brain lags behind. The type of brain injury is dependent on the magnitude of this strain and the time duration over which it acts on the brain.
Risk of focal head and brain injury is measured in terms of peak linear acceleration associated with impact, while risk of diffuse brain injury is measurable in terms of peak angular acceleration.
While helmets can prevent fatalities associated with penetrating head trauma, it may be argued that protection against diffuse brain injury is of paramount importance. After all, cuts, bruises and even bone fractures will heal, but brain injuries often have life long neurologically devastating effects.
Unfortunately, helmet testing standards addresses only the risk of blunt force trauma, not risk of brain injury.
Helmets may reduce the rotational forces acting on the brain. But since helmets are not currently certified according to their ability to protect against brain injury the level of protection is not standardized. Hence, it is possible to sustain catastrophic diffuse brain injuries, even while wearing a helmet.
As a biomechanics researcher, Dr. John Lloyd has dedicated his career to understanding the biomechanics of brain injuries. One objective of which is to develop a new generation of helmets for sports and motorcycling using “intelligent” materials that hold great promise for reducing the risk of traumatic brain injuries.
Dr. Lloyd’s biomechanics laboratory employs a specialized helmet testing apparatus for evaluating the risk of both head and brain injuries. This apparatus has been published in a peer-reviewer journal.
Using this apparatus, Dr. Lloyd, evaluates the linear and rotational forces associated with specific impact events, such as a motorcycle crash or sports injury, to determine whether an unhelmeted condition, or the type of helmet might have prevented the injury sustained. This apparatus has also been used to investigate whether a particular helmet failed to perform or did not meet scientifically-acceptable levels of protection.
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 (13.4 mph). In general, if the resultant 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 motorcycle helmet can provide protection against diffuse brain injuries, including concussion.
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 [i] 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.
Dr. Lloyd recently conducted independent testing of various motorcycle helmets utilizing a methodology that has been peer-reviewed [i] and has survived a Daubert motion for exclusion [ii]. The following figure presents peak angular acceleration results of repeated testing of various motorcycle helmets, including: (i) Voss novelty helmet, (ii) Bell shorty helmet, (iii) Daytona shorty helmet, and (iv) Bell full-face helmet, compared with an unhelmeted condition for impacts onto concrete at approximately 20mph. The red horizontal line on the figure indicates the 50% threshold for concussive trauma, as defined by Pellman et al [iii].
Results show that while a novelty or DOT approved motorcycle helmet will reduce the peak angular acceleration associated with a head impact relative to an unhelmeted condition, the level of protection is not sufficient to prevent diffuse brain injury in a typical motorcycle accident.
[i] Caccese V, Ferguson J, Lloyd J, Edgecomb M, Seidi M and Hajiaghamemar M: Response of an Impact Test Apparatus for Fall Protective Headgear Testing Using a Hybrid-III Head/Neck Assembly. Experimental Techniques, 2014.
[ii] Superior Court, Judicial District of Hartford, CT. Docket Number: HHD-CV-13-6043998-S. Case Caption: SHUMBO, JAKE Et Al v. K2 SPORTS USA Et Al. Order #227.86 regarding: 03/02/2015 Motion to Exclude Expert Testimony. Notice Issued: 07/09/2015
[iii] Pellman EJ, Viano DC, Tucker AM, Casson IR, Waeckerle JF: Concussion in professional football: reconstruction of game impacts and injuries. Neurosurgery 53(4): 799-812, 2003
[iv] COST-327 report of the European Commission Directorate General for Energy and Transport on Motorcycle Safety Helmets. (1999).
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.
football helmet prototype
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.
helmet prototype reduces concussion and brain injury risk
“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 unique capability as a helmet expert is in the biomechanical evaluation of helmets, specifically, football, sports and motorcycle helmets. Helmets are designed to reduce the risk of blunt force trauma to the head, however protection against diffuse traumatic brain injury is often inadequate. Dr. Lloyd is often called upon to opine whether the head and brain injuries may or may not have been prevented by head protection.
Using this apparatus, helmet expert Dr. Lloyd and neurologist Dr. Frank Conidi of the Florida Center for Headache and Sports Neurology have presented a series of studies at the American Academy of Neurology meetings. Their work was publicized by the AAN in a press-release titled “How Well Do Football Helmets Protect Players from Concussions“. Dr. Lloyd and Dr. Conidi published a scientific article in the Journal of Neurosurgery titled “Brain Injury in Sports“. This article documents the limited protection against traumatic brain injuries afforded by many current varsity football helmets.
Example Motorcycle Helmet Expert Case
Dr. Lloyd provided biomechanical analysis on a recent motorcycle accident case in which an automobile crossed the path of an unhelmeted rider traveling at approximately 25 miles per hour. The motorcyclist’s head shattered the driver’s side window, leading to catastrophic brain injury
Dr. Lloyd was asked to opine as to whether or not a motorcycle helmet would have prevented these injuries. A test apparatus was constructed using an exemplar automobile driver’s door and window to measure the forces acting on a crash test dummy head.
The following high speed videos and data were captured showing the helmeted versus unhelmeted conditions.
We learned that, had the motorcyclist been wearing a helmet during the subject collision, he would have most likely sustained fatal neck injuries as the helmet was deflected by the window, producing unsurvivable neck extension, as shown below:
Test results and video documentation were presented at deposition and proved highly valuable.
Please call Dr. Lloyd at 813-624-8986 or email DrJohnLloyd@Tampabay.RR.com to discuss how he can be of help to you with your case.
The humble helmet dates back nearly 3000 years and though it has been used prolifically in warfare, it is now most commonly used to provide head protection outside the combat arena.
However, although applications might have diversified, it is still fundamentally designed and used to provide the same thing.
So when this most traditional of objects is combined with modern sensor technologies, greater test data resolution and analysis, there is bound to be fresh insight.
And this is the case for many conventional designs where sensors, test and measurement technologies are changing conventional thinking into how something has been designed, to how it should be designed.
It sets the scene and means helmet design is on a collision course for further impact protection, specifically in preventing serious brain injury by giving helmet designers greater clarity in to the mechanical forces at play in any particular scenario.
It was this, along with a lifetime of comprehensive knowledge, which enabled biomechanist Dr John Lloyd, research director of BRAINS, to start up a company dedicated to improving current helmet technology and ultimately improve protection for wearers. He aims to shed new light on helmet design, and improve protection against the fundamental causes of concussion and brain injury.
“There are two key forces at play during a head impact,” said Dr Lloyd, speaking at this year’s National Instruments Week in Austin, Texas. “Firstly there are linear forces, these are the ones that cause visible injuries such as bruising and skull fractures. However, the second is the rotational forces. These are the ones that cause invisible injuries such as concussion and brain injury.
“Current helmet testing technologies measure the linear forces. However, at this time, they do not measure the rotational forces, so consequently we have helmets for many sports that do not test against their ability to provide protection against concussions and brain injury.”
Whether it is for riding a bike, horse riding, skiing or indeed for the soldier in the field, the effect of rotational movement is the same. Yet, it is rarely tested for, and even less frequently measured, to see how effective any helmet is in rotation force protection.
Dr Lloyd modified the standard apparatus used for testing helmets (see the rig on page 28), where a head section is raised 2m on a rig and dropped under gravity before it hits a striking plate with an impact force in the region of 4500N. However, instead of using a standard head form, Dr Lloyd replaced it with a standard automotive crash test dummy head and neck section. This way, when the head impacts the striking plate at the bottom of the test rig it will rotate, and the movement measured.
“We had multiple sensors embedded in the centre of mass of this head form,” explained Lloyd. “So, during the impact we were able to measure the linear acceleration as well as the angular motion of the head.
“My measuring apparatus includes sensors from several manufacturers.. The angular rate sensor, for example, that is used to measure the rotational forces is a highly specialised sensor. And, as a result, has its own data acquisition hardware and software.”
Simplifying synchronisation
Trying to integrate all this data from different sensors was a challenge at best. And to make matters more complicated, the peak linear acceleration and peak angular acceleration actually happen at different points in time.
“So while you can just line up the data,” he said, “there is a lag between them. So we need to measure that lag, which is a critical measurement in the research.”
To resolve the problem, Dr Lloyd uses both the National Instruments LabView graphical software and a CompactDAQ to interface with the sensors and provide the necessary synchronisation between the various sensors.
Dr Lloyd modified his apparatus for testing helmets used by American footballers in the National Football League (NFL), to develop understanding of the how spinal and head injuries are caused and improve the design of the standard helmet.
“The results are pretty alarming in terms of how little protection they provide against concussions and traumatic brain injuries,” he said.
“Based on lessons learned from that study, I have developed a new ‘football’ helmet prototype. This uses a patent pending matrix of non-Newtonian materials and when we tested the prototype helmet, on the same apparatus, the result blew me away. Not only did these materials reduce the linear forces but compared to the standard football helmet they actually reduced the rotational forces that cause concussion and brain injury by an amazing 50%.”
The non-Newtonian materials Lloyd has in mind are inexpensive and produce a helmet that is considerably lighter and even said to be more comfortable for those wearing them.
Dr Lloyd is now expanding the concept of reducing rotation forces in helmets in every application and said it can be applied to almost any helmet design to help reduce concussion and brain injuries from sports to leisure and even back to warfare.
Building a rig and conducting the test
A modification to the US National Operating Committee on Standards for Athletic Equipment (NOCSAE) standard test apparatus was used by Dr John Lloyd, research director of US helmet research start-up, BRAINS.
He developed and validated a new helmet test rig to measure the impact of protective headwear to include measurements of both linear and angular kinematics. This apparatus consists of a twin wire fall test system equipped with a drop arm that incorporates a 50th percentile Hybrid III head and neck assembly from HumaneticsATD crash test dummy, as used in the automotive industry.
The aluminium fly arm runs on Teflon sleeves through parallel braided stainless steel wires, which are attached to mounting points in the building structure and anchored into the concrete foundation. The anvil, onto which the head drop systems impacts, consists of a 350mm x 350mm steel based plate.
Both the standard Riddell Revolution Speed US university football helmet, and the prototype BRAINS helmet that incorporates a non-Newtonian matrix, were dropped from a height of 2m onto a flat steel anvil, in accordance with American Society for Testing and Materials (ASTM) standards. This generated an impact velocity of 6.2 m/s (13.9 mph).
Instrumentation:
A triaxial accelerometer from PCB Piezotronics and three DTS-ARS Pro 18k angular rate sensors (Diversified Technical Systems) were affixed to a tri-axial block installed at the centre of mass in the Hybrid III head form. Data from the accelerometer and angular rate sensors were acquired using National Instruments compactDAQ hardware.
Analysis:
Data from the analogue sensors were acquired at 10,000Hz, per channel, using LabView and then filtered in Matlab using a phaseless 4th order Butterworth filter with a cut off frequency of 1650Hz. Angular acceleration values were derived from the angular velocity data based on a 5-point least squares quartic equation.
Result:
The result of the new helmet design shows significant improvement in rotational acceleration exerted on the head and neck, cutting the overall force by nearly 50%.
Author
Justin Cunningham
– See more at: http://www.eurekamagazine.co.uk/design-engineering-features/technology/why-all-head-protection-is-in-need-of-a-redesign/66493/#sthash.6Tv5duXE.dpuf
The common belief among riders is that a motorcycle helmet protects the whole head, including the brain. However testing standards in Europe (ECE 22.05) and the US (DOT & Snell), which involve dropping helmeted headforms from heights of 2-3 meters onto a steel plate, only evaluate a motorcycle helmet in terms of its ability to protect against blunt force trauma, such as skull fractures and penetrating head injuries. The mechanism underlying diffuse brain injuries, such as concussions and brain hemorrhages is distinctly different, but is not assessed by current motorcycle helmet testing standards.
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 
By design, helmets reduce impact force by increasing the impact duration. But, as demonstrated by 
