Accident Analysis and Prevention 70 (2014) 1–7
Contents lists available at ScienceDirect
Accident Analysis and Prevention journal homepage: www.elsevier.com/locate/aap
Bicycle helmets are highly effective at preventing head injury during head impact: Head-form accelerations and injury criteria for helmeted and unhelmeted impacts Peter A. Cripton a,b,c,d,∗ , Daniel M. Dressler a,b,d , Cameron A. Stuart e , Christopher R. Dennison a,b , Darrin Richards e a
Department of Mechanical Engineering, University of British Columbia, Vancouver, Canada International Collaboration on Repair Discoveries, University of British Columbia, Canada c Centre for Hip Health and Mobility, University of British Columbia, Canada d Orthopaedic and Injury Biomechanics Group, University of British Columbia, Canada e Synaptic Analysis Consulting Group, Vancouver, British Columbia, Canada b
a r t i c l e
i n f o
Article history: Received 18 February 2013 Received in revised form 8 January 2014 Accepted 19 February 2014 Keywords: Brain injury Concussion Helmet Bicycle Injury prevention
a b s t r a c t Cycling is a popular form of recreation and method of commuting with clear health benefits. However, cycling is not without risk. In Canada, cycling injuries are more common than in any other summer sport; and according to the US National Highway and Traffic Safety Administration, 52,000 cyclists were injured in the US in 2010. Head injuries account for approximately two-thirds of hospital admissions and threequarters of fatal injuries among injured cyclists. In many jurisdictions and across all age levels, helmets have been adopted to mitigate risk of serious head injuries among cyclists and the majority of epidemiological literature suggests that helmets effectively reduce risk of injury. Critics have raised questions over the actual efficacy of helmets by pointing to weaknesses in existing helmet epidemiology including selection bias and lack of appropriate control for the type of impact sustained by the cyclist and the severity of the head impact. These criticisms demonstrate the difficulty in conducting epidemiology studies that will be regarded as definitive and the need for complementary biomechanical studies where confounding factors can be adequately controlled. In the bicycle helmet context, there is a paucity of biomechanical data comparing helmeted to unhelmeted head impacts and, to our knowledge, there is no data of this type available with contemporary helmets. In this research, our objective was to perform biomechanical testing of paired helmeted and unhelmeted head impacts using a validated anthropomorphic test headform and a range of drop heights between 0.5 m and 3.0 m, while measuring headform acceleration and Head Injury Criterion (HIC). In the 2 m (6.3 m/s) drops, the middle of our drop height range, the helmet reduced peak accelerations from 824 g (unhelmeted) to 181 g (helmeted) and HIC was reduced from 9667 (unhelmeted) to 1250 (helmeted). At realistic impact speeds of 5.4 m/s (1.5 m drop) and 6.3 m/s (2.0 m drop), bicycle helmets changed the probability of severe brain injury from extremely likely (99.9% risk at both 5.4 and 6.3 m/s) to unlikely (9.3% and 30.6% risk at 1.5 m and 2.0 m drops respectively). These biomechanical results for acceleration and HIC, and the corresponding results for reduced risk of severe brain injury show that contemporary bicycle helmets are highly effective at reducing head injury metrics and the risk for severe brain injury in head impacts characteristic of bicycle crashes. © 2014 Elsevier Ltd. All rights reserved.
1. Introduction
∗ Corresponding author at: Department of Mechanical Engineering, University of British Columbia, 6250 Applied Science Lane, Vancouver, BC V6T 1Z4, Canada. Tel.: +1 604 675 8835. E-mail address: [email protected] (P.A. Cripton). http://dx.doi.org/10.1016/j.aap.2014.02.016 0001-4575/© 2014 Elsevier Ltd. All rights reserved.
Cycling is a popular form of recreation and it is used for commuting and other forms of transportation. It is generally safe and the health benefits of it are clear (Hamer and Chida, 2008; Wen and Rissel, 2008), which is in sharp contrast to motorized transportation of any type. However, cycling is also not without risk. In Canada, cycling injuries are the most common injury occurring from summer sports; over 4300 people were hospitalized due to a cycling
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injury in 2009–2010 (Canadian Institute for Health Information, 2010). According to the National Highway and Traffic Safety Administration (NHTSA), between 600 and 800 cyclists are fatally injured each year in the United States and 52,000 cyclists were injured in the US in 2010 (NHTSA Traffic Safety Facts 2010 Data, 2010). Among cyclists, head injuries account for approximately two-thirds of hospital admissions and three-quarters of fatal injuries (Thompson et al., 1999). Epidemiological studies show that helmets are highly effective at preventing head and brain injury amongst riders who crash. A case–control study conducted by Thompson et al. (1989) in Seattle over a period of 1 year found that bicycle helmets reduced the risk of head and brain injury by 85% and 88%, respectively. In a second larger case–control study by the same group (Thompson et al., 1996), helmets decreased the risk of head injury by 69%, brain injury by 65%, and severe brain injury by 74%. Helmets were found to be equally effective in accidents involving motor vehicles and those not involving motor vehicles. Furthermore, helmets were found to provide substantial protection from head injuries across all age groups. Amoros et al. (2012) recently conducted a case–control study in France and studied helmet effectiveness over more than 13,500 cyclist injuries. They concluded that helmets were associated with a decreased risk of head injury in cyclist trauma and this decrease seemed to be more pronounced for severe head injuries. Maimaris et al. (1994) studied over a thousand patients that sustained cycling-related injuries who were treated at an emergency department in England. They concluded that helmets reduced the risk of head injury by a factor of more than three. Heng et al. (2006) found that helmet use significantly reduced the risk of head and facial injury in a 2006 study of cycling trauma in Singapore. Despite the protection provided by helmets, as demonstrated by the epidemiological studies above, the safety benefits offered by helmets are not universally accepted. Many cities, towns, states and provinces do not have helmet laws and many cyclists do not wear helmets (Page et al., 2012). Anti-helmet groups state that helmets are not effective and that, in some cases, due to the increased size of a helmeted head compared to a bare head or due to the compliance of the shell or presence of vent holes, helmets can cause “rotational” injuries such as diffuse axonal injury (DAI). In the lay press, some groups claim that helmets cause injuries by obstructing vision or blocking sound. Researchers have also published articles, critical of the many epidemiological studies (cited above) that show that helmets are highly effective at preventing head injuries, accusing them of bias and conflicts of interest (Curnow, 2006; Elvik, 2011). Curnow argued that bicycle helmets are not as effective as claimed because previous epidemiological studies have not considered rotational injury (Curnow, 2003). There is considerable debate on the merit and limitations of the epidemiological evidence (Curnow, 2006, 2003; Elvik, 2011; Hagel and Barry Pless, 2006). One limitation of the epidemiological approaches is that it infers helmet performance during the impact from evidence collected after the impact and thus the severity of the head impact under study is never known. It is not our purpose to debate the merit of the epidemiological literature. Here we aim to explore the extent to which the epidemiological evidence of helmet efficacy can be supported or contradicted by
a biomechanical study that allows study of helmet performance during the impact. Biomechanical investigations of helmet efficacy, and indeed helmet certification standards, simulate helmeted head impact by dropping helmeted headforms onto prescribed impact surfaces. In helmet certification standards, the primary metric to assess impact management efficacy is linear headform acceleration measured during a drop test; helmets are considered to have met the certification criteria if the helmeted headform acceleration is below a prescribed threshold. The threshold varies from standard to standard (Table 1), and is not directly correlated to established risk curves. The standards generally require that helmets be certified using a magnesium headform. The range of drop heights associated with these standards is from 1.5 m (EN1078) to 2.2 m (Snell B95A) (Table 1). In biomechanical investigations, linear and rotational head accelerations are measured during the impact and helmet efficacy is determined by comparing these accelerations, and other derived metrics such as the Head Injury Criterion (HIC), to injury risk functions. For example, Mertz et al. have established head injury probability curves, in terms of HIC and linear acceleration, for the Hybrid III headform which was originally developed for automotive crash testing (Mertz et al., 2003). Because injury tolerances exist for this headform, the Hybrid III is increasingly applied in biomechanical helmet and head impact studies (Beckwith et al., 2012; Kendall et al., 2012; Pang et al., 2011; Pellman et al., 2003; Scher, 2006; Scher et al., 2009; Viano and Halstead, 2012; Viano and Pellman, 2005). Overall, the biomechanical studies indicate that helmets significantly reduce head accelerations relative to unhelmeted impacts (Benz et al., 1993; Hodgson, 1990; Mattei et al., 2012; Scher, 2006) or to impacts with thin uncertified novelty helmets (DeMarco et al., 2010; Scher et al., 2009). Furthermore, because linear head acceleration is known to be monotonically correlated to concussion and skull fracture risk (Greenwald et al., 2008; Mertz et al., 2003; Pellman et al., 2003) they are therefore known to reduce the risk of sustaining head injury. The biomechanical comparison that best matches the epidemiological studies, and thus that would be best able to augment the debate in that field, is a comparison of helmeted and unhelmeted head impact under identical impact conditions. Unfortunately, these tests are difficult to perform because of limitations of the magnesium head forms that are mandated in bicycle helmet standards and that have thus most often been used in bicycle helmet impact tests. The magnesium head forms are at high risk of damage if they are tested with no helmet and they have not been validated to match the expected human response for bare head impacts. Thus there is no test series available to our knowledge that contrasts helmeted and unhelmeted impacts for contemporary bicycle helmets under direct matched impact. Hodgson contrasted early 1990s era helmets with an unhelmeted impact using a “small humanoid headform”, Benz et al. dropped an unhelmeted Hybrid II headform from a lower height than their helmeted impacts and Mattei et al. dropped human cadaver skulls with and without helmets from six and nine inch drop heights (Benz et al., 1993; Hodgson, 1990; Mattei et al., 2012). All of these studies demonstrated a dramatic decrease in head accelerations for the helmeted compared to the
Table 1 Comparison of several bicycle helmet standards. Standard
Reference
Drop height (m)
Drop height (feet)
Criteria (g’s)
Consumer Product Safety Commission (CPSC) Snell Memorial Foundation (Snell) American Society for Testing and Materials (ASTM) Canadian Standards Association (CSA) European Standards (CEN)
16 CFR Part 1203 BF95 (1998 Revision) ASTM F1447F12 CSA D113 2FM89 (Reaffirmed 2004) EN 1078
2 2.2 2 1.6 1.5
6.6 7.2 6.6 5.2 4.9
300 300 300 250 250
P.A. Cripton et al. / Accident Analysis and Prevention 70 (2014) 1–7
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Fig. 1. Photograph showing helmeted Hybrid III headform (left) and unhelmeted Hybrid III headform (right) in contact with the steel anvil. The Hybrid III headform was attached to the ball arm which was mounted to a linear bearing on the monorail drop-tower.
unhelmeted impacts. However, to the best of our knowledge, there have been no studies that compare the crucial situation of helmeted and unhelmeted impacts using contemporary bicycle helmets, with a head form that is validated for bare head impacts and from drop heights that compare to bicycle helmet standards and real world cycling falls. The data that would result from such a test series is directly relevant and indeed central to the ongoing debate of bicycle helmet efficacy. The objective of this study was to assess the biomechanical efficacy of bicycle helmets to reduce risk of head injury in simulated head impacts from drop heights consistent with bicycle helmet standards and real world cycling head impacts. We designed and fabricated a custom-made test fixture that allowed us to attach a Hybrid III headform to a monorail drop tower and performed both helmeted and unhelmeted drops. The Hybrid III head form can be tested without a helmet and it is validated in bare head impacts (Foster et al., 1977). Linear head acceleration was measured, and HIC and injury risk were determined from these accelerations to ascertain the efficacy of helmets to reduce risk of head injury. 2. Methods and materials We simulated head impacts using a monorail drop tower similar to those specified in helmet certification standards. We fabricated a custom-made test fixture that allowed us to attach a Hybrid III headform (Humanetics Inc., Plymouth, MI, USA), that corresponded to a 50th percentile male head, to the drop tower. A ball-arm was mounted to a monorail drop tower that was purpose-built for this application. Paired tests were performed in order to study the risk of injury in helmeted and unhelmeted impacts. A paired test is defined as two drops onto an identical anvil and from identical drop heights both with and without a helmet. The impact surface for all drops was a flat, fixed steel anvil. Fig. 1 shows the anvil, helmeted and unhelmeted Hybrid III headform and features of the bearing and guide rail. Translational acceleration along the direction of impact was measured using a single axis accelerometer (±2000 g range, Endevco model 7264C2000, Meggitt Sensing Systems, San Juan Capistrano, CA, USA), which was mounted to the center of the ball-arm which in turn was placed within the Hybrid III head close to the head center of mass. The mass of the entire drop assembly, including the Hybrid III headform, ball-arm, and linear bearing was 5.05 kg. The mass of the helmet was approximately 0.25 kg and was considered additional mass for the helmeted drops. The helmeted and unhelmeted drops were conducted from nominal heights starting at 0.5 m to 3 m in 0.5 m increments. This range brackets heights used in certification standards but exceeds
the maximum height of typical standards (CSA D113.2-M89 (1.7 m), CPSC (2.0 m), ASTM F1447 (2.0 m), EN1078 (1.5 m) Snell B95A (2.2 m)) to allow study of higher energy impacts that can occur in real-world cycling where falls can happen while traveling at considerable speed. Our testing range also brackets the range of perpendicular impacts documented for reconstructed bicycle falls (Fahlstedt et al., 2012). Two drops were performed at 0.5 m, 1 m, 1.5 m, 2.5 m and 3 m; one drop for a helmeted Hybrid III headform and one unhelmeted. Six drops, three helmeted and three unhelmeted, were performed from 2 m. More drops were performed at 2 m than other heights to obtain drop speed and acceleration data that would allow a limited investigation of the repeatability of the experiment and to do so at common drop height used in bicycle helmet standards. To assess repeatability, we calculated the maximum difference in both peak acceleration and HIC and expressed these differences as a percentage of the mean peak acceleration and HIC. In total sixteen drops were conducted (8 helmeted, 8 unhelmeted). The headform was adjusted so that impacts took place to the forehead of the headform as seen in Fig. 1. Actual drop heights were increased by approximately 5 cm above the nominal drop height to account for friction in the drop rail. Speed at impact was calculated using high-speed video and was found to be within 5% of the expected velocity for each respective drop height. All helmets used in this work were CCM V15 Backtrail bicycle helmets (Reebok-CCM Hockey, Montreal, QC, Canada). The helmets were constructed with a micro-shell and an expanded polystyrene liner. The helmets conformed to the standards set out by the Consumer Product Safety Commission (CPSC) (CPSC, 1998). In impacts where helmets were used, the helmet was placed on the Hybrid III headform in a standardized fashion. The orientation of the headform was held constant for all drop heights using angle and position landmarks drawn on the Hybrid III headform (Fig. 1). The chin retention strap was tightened to secure the helmet to the Hybrid III headform (Fig. 1). The helmets were also equipped with a ratcheting tension system that is designed to pass inferior to the occipital protuberance. This was tightened prior to all drop tests. A check for helmet fit and secure attachment to the Hybrid III headform involved manipulation of the helmet on the head to ensure no visible relative motion. Each helmet was used for a single drop and then replaced with a new helmet. An Analog Devices (Analog Devices Inc., Norwood, MA) data acquisition system was used to collect the data, with the acceleration signal sampled at 39 kHz and hardware anti-alias filtered to comply with SAE J211-1 (“SAE J211 Instrumentation for Impact Test – Part 1: Electronic Instrumentation”). In addition, the accelerometer data were low-pass filtered at 1650 Hz (CFC1000) during post-processing as per SAE J211-1.
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Fig. 2. Typical acceleration data plotted versus time for both a helmeted and unhelmeted Hybrid III. Data shown is for the 2 m drop height. Acceleration is expressed in g, where one g corresponds to 9.81 m/s2 .
Peak head accelerations were used as a biomechanical metric of helmet efficacy and were compared for the paired helmeted and unhelmeted tests. In order to assess the risk of injury associated with these impacts, head accelerations were also compared to the Injury Assessment Reference Value (IARV), and probability curves published by Mertz et al. (2003). For presenting peak acceleration data, we use a 5% risk threshold for skull fracture based on peak acceleration (180 g). The Head Injury Criterion (hereafter HIC) was calculated during post-processing using Eq. (1). The HIC quantifies head impact severity by incorporating time of acceleration exposure and acceleration magnitude.
HIC15 =
1 t2 − t1
2.5
t2
a(t)dt t1
(t2 − t1 )
(1) max
For this analysis, a(t) is the head acceleration, in g, as measured by the single axis accelerometer, and the time interval (t2 − t1 ) was chosen to maximize HIC over a maximum duration of 15 ms (Eppinger et al., 1999). In the subject testing the sensing axis of the accelerometer was aligned with the direction of impact and thus captured the resultant acceleration. Similar to the acceleration analysis, HIC15 values were compared to the IARV of 700 which Mertz et al. have reported corresponds to a 5% risk of AIS ≥ 4 brain injury for the adult population (Mertz et al., 2003).
Fig. 3. Peak accelerations for both helmeted and unhelmeted drops. Numbers over bars indicate peak acceleration. For 2 m drop height, results stated are the mean value calculated from three drops. Horizontal dashed line indicates the IARVof 180 g (5% chance of skull fracture) for a midsize male.
probability were all of smaller magnitude in drops where the Hybrid III was helmeted. The duration of the impact pulse was larger in helmeted drops relative to unhelmeted. As the head decelerated, a small amount (approximately less than 5 mm) of sliding outward occurred between the helmet shell and the impact surface as the helmet shell and liner deformed during impact. Peak accelerations (Fig. 3) were smaller in helmeted drops relative to unhelmeted drops, for all drop heights. On average and considering all drops, the peak accelerations for helmeted drops were smaller by a factor of 4.2 relative to unhelmeted. For the severity of impacts tested, peak acceleration exhibited a linear relationship with drop height. In the unhelmeted situation, the head accelerations were above the IARV of 180 g for every drop from 0.5 m to 3 m. For drop heights of 0.5–1.5 m, helmets decreased the peak accelerations to a value below the IARV (Fig. 3). Fig. 4 shows maximum HIC for helmeted and unhelmeted drops from all heights. For each drop height, helmets reduced HIC relative to unhelmeted drops. The mean interval required to maximize HIC for unhelmeted and helmeted drops was 1.0 ms and 5.0 ms, respectively. The increased HIC interval for helmeted drops is consistent with the considerably wider (in the time domain) acceleration peak for helmeted drops relative to unhelmeted shown in Fig. 2. Fig. 5 shows the calculated risk of a severe brain injury (AIS 4+) for all helmeted and unhelmeted drops. Overall, the helmeted drops dramatically reduced the risk across all drop heights. For drops of 1 meter and greater, the unhelmeted condition resulted in essentially
3. Results Repeatability was evaluated by analyzing multiple drops at 2 m (3 helmeted drops and 3 unhelmeted drops). The maximum interdrop difference in peak acceleration was 1.5% and 3.3% (percentage of mean peak accelerations for 2 m drop) for the helmeted and unhelmeted Hybrid III headform, respectively. Similarly, maximum inter-drop differences in HIC were 6.0% and 5.0% (percentage of mean HIC), respectively. Fig. 2 shows typical acceleration curves plotted over the time of the impact event for both the unhelmeted and helmeted Hybrid III headform. In general, the acceleration magnitudes plotted over time exhibited a single abrupt increase in acceleration, which continues to the peak acceleration, followed by an abrupt decrease in acceleration. Following this, the accelerations fluctuate (e.g. Fig. 2, after 5 ms for the unhelmeted data) and these fluctuations correspond to head/helmet–anvil interactions that are secondary to the initial head/helmet-to-anvil impact (i.e. the head “bounces” off of the anvil). In general, peak accelerations, HIC and injury
Fig. 4. Head Injury Criterion (HIC) calculated using HIC15 convention for both helmeted and unhelmeted drops. Numeric values over bars indicate HIC values and long dashed line indicates IARV based on HIC.
P.A. Cripton et al. / Accident Analysis and Prevention 70 (2014) 1–7
Fig. 5. Calculated probability of sustaining a severe brain injury (i.e. AIS 4+) based on the HIC values for the helmeted and unhelmeted drops. Numeric values over bars indicate risk magnitude.
100% chance of an AIS 4+ brain injury. For all drops of 2 m or less, the helmeted condition resulted in a risk of below 35%. However, at the 3 m drop height the risk of an AIS 4+ brain injury exceeded 90% for the helmeted condition. For each of the helmeted drops, various levels of helmet deformation and helmet shell and foam damage were evident. The expanded polystyrene (EPS) foam cracked near the impact location in all but the lowest drop height (0.5 m) and the micro-shell fractured in the 3 m drop. Upon further inspection, all of the test scenarios resulted in plastic deformation of the EPS foam liner near the impact location. 4. Discussion The overarching objective of this study was to assess the efficacy of certified contemporary bicycle helmets to mitigate skull and brain injury risk in head impacts with characteristic velocities matching impact velocities from helmet certification standards and also matching those from the biomechanical literature. We used a purpose-built drop tower with a Hybrid III test headform and a uniaxial accelerometer aligned to the direction of impact to measure linear head accelerations during both helmeted and unhelmeted drops. Linear head acceleration is one accepted mechanical measure that can be related to both skull injury risk and brain injury risk (Mertz et al., 2003; Pellman et al., 2003) both directly and through HIC (Mertz et al., 2003), and therefore acceleration was our primary biomechanical measure to assess efficacy. We have characterized the ability of one typical contemporary bicycle helmet to reduce the severity of a head impact and reduce the risk of severe life-threatening skull and brain injury, compared to not wearing a helmet, in matched impact tests where impact severities (i.e. drop height and pre-impact head velocity) were identical for the case of the helmeted and unhelmeted headform. The tested helmet dramatically decreased peak linear head acceleration (Fig. 3), HIC15 (Fig. 4) and the potential for severe brain injury (Fig. 5) in all impacts. Considering peak acceleration (Fig. 3), the helmeted headform experienced accelerations below the IARV of 180 g (Mertz et al., 2003) in drops from 0.5 m, 1.0 m, and 1.5 m while the unhelmeted headform experienced acceleration well above the IARV in drops from all heights. In drops from 2.0 m up to 3.0 m, the helmeted headform experienced accelerations above the IARV, but helmeted headform accelerations were at least 4 times smaller than those of the unhelmeted headform (Fig. 3). Evaluation of HIC15 (Fig. 4) was consistent with the acceleration results when compared with the IARV of 700. Probability of skull fracture (not shown)
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and severe brain injury (Fig. 5) is reduced, for all drop heights, for the helmeted headform relative to the unhelmeted headform. This biomechanical evidence clearly indicates that contemporary bike helmets are highly effective at reducing injury risk through paired helmeted and unhelmeted impacts with realistic drop heights and impact speeds. For example, the helmets reduced the head peak acceleration from 824 g to 181 g for drops of 2.0 m reducing the risk of skull fracture from 99.9%+ to 5%. The reductions in head acceleration and HIC15 described above demonstrate that certified helmets significantly reduce the risk of sustaining severe and even fatal injuries. It is worth noting that common helmet standards do not seem to be designed to prevent head accelerations from exceeding the IARV values for severe skull and brain injuries published by Mertz et al. (2003), although it is noted that the helmets reduced the accelerations to a value well below that specified in the standard. The helmeted drop from 2.0 m resulted in a HIC of 1250 that corresponds to a 34% chance of severe brain injury. It may be necessary to re-evaluate the helmet standards and contemplate lowering the allowable accelerations in some standards and to require testing from multiple drop heights in order to decrease the potential for serious and severe skull and brain injuries from falls of various drop heights. Common injuries coded AIS 4 (severe) and above include penetrating skull injuries leading to brain injury, large contusions (e.g., coupe-contrecoup, intermediate, and gliding) and hematomas (e.g., subdural, subarachnoid, and intracerebral), as well as diffuse axonal injury with associated loss of consciousness for a period exceeding six hours. The severity of these injuries when coded as 4+ range from severe to maximum (usually fatal) (Gennarelli and Wodzin, 2005). Considering a realistic bicycle accident scenario documented in the literature (Fahlstedt et al., 2012) where a cyclist was thrown at 20 km/h (i.e. 5.6 m/s which corresponds to a drop height of approximately 1.5 m), our analysis indicates that a helmeted cyclist in this situation would have a 9% chance of sustaining the severe brain and skull injuries noted above whereas an unhelmeted cyclist would have sustained these injuries with 99.9% certainty. In other words, a helmet would have reduced the probability of skull fracture or life threatening brain injury from very likely to highly unlikely. Evaluation of the 3 m drops demonstrate that helmets only offer a finite amount of protection. However, the 7.7 m/s impact speed is not representative of most real-world bicycle impacts (Fahlstedt et al., 2012). At impact speeds of this velocity the energy management capability of the helmet is saturated (Newman, 2002) and the EPS liner bottoms out. The range of drop heights that we tested was consistent with the range of impact speeds that has been documented as plausible for cyclist impact scenarios. Our 3.0 m, 2.5 m, and 2.0 m drops were consistent with the resultant cyclist head velocity in the Fahlstedt study that was oblique to the ground (i.e. glancing off it). However, if the cyclist hit a curb or the wheel of a car at the point of glancing off the ground then there would be head impact of similar severity and impact direction to the higher impact severity tests that we carried out. The biomechanical results for acceleration and HIC15 , and the corresponding results for reduced risk of severe skull and brain injury, complement epidemiological studies that have sought to assess the protective efficacy of bicycle helmets. The majority of the epidemiological studies suggest that helmets are effective at reducing injury risk in a range of sports and across both adult and youth segments of the population. However, concerns have been raised in recent years over issues of selection-bias, failure to compensate for time-trending and public policy, and lack of control for confounding aspects of the input data for statistical studies including type and mechanism of head/brain injury and severity of the head impact (Curnow, 2006; Elvik, 2011). A fundamental reason for this study is that the lack of appropriate knowledge of and
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statistical control for the severity of the impact confounds the results and decreases the significance of results below what they would be if appropriate controls were applied. As a consequence, the results in some studies suggest helmets are only slightly effective at preventing head injuries or, in extreme cases, actually cause head injuries (Elvik, 2011). However, the results of our biomechanical study, where impact severity was controlled in helmeted and unhelmeted impacts, strongly refute the epidemiological studies that suggest helmets are marginally effective. Indeed, in all impacts the risk of sustaining injury was reduced when the Hybrid III headform was helmeted, and in the case of a realistic bicycle head impact (5.5 m/s impact speed which corresponds to a drop height of approximately 1.5 m), a helmeted cyclist would have a 9% chance of sustaining AIS 4+ injuries whereas an unhelmeted cyclist would almost certainly sustain these injuries. The results of this study are in good agreement with previous biomechanical research on the protective efficacy of bicycle helmets. For example, Benz et al. (1993) conducted a study of child and adolescent helmets in drops from 1 m to 1.5 m using an unspecified “child headform” (Benz et al., 1993) and a Hybrid II headform for adolescent helmets, both dropped onto a flat anvil, and reported an overall threefold decrease in HIC as a result of protecting the headform with a helmet. Hodgson (1990) conducted a study using a Hodgson-WSU headform dropped from 1 m and 2 m onto flat and convex surfaces (Hodgson, 1990). For 1 m drops, the unhelmeted headform was at increased risk of injury (70–99% of the population would be injured) relative to the helmeted headform (
Contents lists available at ScienceDirect
Accident Analysis and Prevention journal homepage: www.elsevier.com/locate/aap
Bicycle helmets are highly effective at preventing head injury during head impact: Head-form accelerations and injury criteria for helmeted and unhelmeted impacts Peter A. Cripton a,b,c,d,∗ , Daniel M. Dressler a,b,d , Cameron A. Stuart e , Christopher R. Dennison a,b , Darrin Richards e a
Department of Mechanical Engineering, University of British Columbia, Vancouver, Canada International Collaboration on Repair Discoveries, University of British Columbia, Canada c Centre for Hip Health and Mobility, University of British Columbia, Canada d Orthopaedic and Injury Biomechanics Group, University of British Columbia, Canada e Synaptic Analysis Consulting Group, Vancouver, British Columbia, Canada b
a r t i c l e
i n f o
Article history: Received 18 February 2013 Received in revised form 8 January 2014 Accepted 19 February 2014 Keywords: Brain injury Concussion Helmet Bicycle Injury prevention
a b s t r a c t Cycling is a popular form of recreation and method of commuting with clear health benefits. However, cycling is not without risk. In Canada, cycling injuries are more common than in any other summer sport; and according to the US National Highway and Traffic Safety Administration, 52,000 cyclists were injured in the US in 2010. Head injuries account for approximately two-thirds of hospital admissions and threequarters of fatal injuries among injured cyclists. In many jurisdictions and across all age levels, helmets have been adopted to mitigate risk of serious head injuries among cyclists and the majority of epidemiological literature suggests that helmets effectively reduce risk of injury. Critics have raised questions over the actual efficacy of helmets by pointing to weaknesses in existing helmet epidemiology including selection bias and lack of appropriate control for the type of impact sustained by the cyclist and the severity of the head impact. These criticisms demonstrate the difficulty in conducting epidemiology studies that will be regarded as definitive and the need for complementary biomechanical studies where confounding factors can be adequately controlled. In the bicycle helmet context, there is a paucity of biomechanical data comparing helmeted to unhelmeted head impacts and, to our knowledge, there is no data of this type available with contemporary helmets. In this research, our objective was to perform biomechanical testing of paired helmeted and unhelmeted head impacts using a validated anthropomorphic test headform and a range of drop heights between 0.5 m and 3.0 m, while measuring headform acceleration and Head Injury Criterion (HIC). In the 2 m (6.3 m/s) drops, the middle of our drop height range, the helmet reduced peak accelerations from 824 g (unhelmeted) to 181 g (helmeted) and HIC was reduced from 9667 (unhelmeted) to 1250 (helmeted). At realistic impact speeds of 5.4 m/s (1.5 m drop) and 6.3 m/s (2.0 m drop), bicycle helmets changed the probability of severe brain injury from extremely likely (99.9% risk at both 5.4 and 6.3 m/s) to unlikely (9.3% and 30.6% risk at 1.5 m and 2.0 m drops respectively). These biomechanical results for acceleration and HIC, and the corresponding results for reduced risk of severe brain injury show that contemporary bicycle helmets are highly effective at reducing head injury metrics and the risk for severe brain injury in head impacts characteristic of bicycle crashes. © 2014 Elsevier Ltd. All rights reserved.
1. Introduction
∗ Corresponding author at: Department of Mechanical Engineering, University of British Columbia, 6250 Applied Science Lane, Vancouver, BC V6T 1Z4, Canada. Tel.: +1 604 675 8835. E-mail address: [email protected] (P.A. Cripton). http://dx.doi.org/10.1016/j.aap.2014.02.016 0001-4575/© 2014 Elsevier Ltd. All rights reserved.
Cycling is a popular form of recreation and it is used for commuting and other forms of transportation. It is generally safe and the health benefits of it are clear (Hamer and Chida, 2008; Wen and Rissel, 2008), which is in sharp contrast to motorized transportation of any type. However, cycling is also not without risk. In Canada, cycling injuries are the most common injury occurring from summer sports; over 4300 people were hospitalized due to a cycling
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injury in 2009–2010 (Canadian Institute for Health Information, 2010). According to the National Highway and Traffic Safety Administration (NHTSA), between 600 and 800 cyclists are fatally injured each year in the United States and 52,000 cyclists were injured in the US in 2010 (NHTSA Traffic Safety Facts 2010 Data, 2010). Among cyclists, head injuries account for approximately two-thirds of hospital admissions and three-quarters of fatal injuries (Thompson et al., 1999). Epidemiological studies show that helmets are highly effective at preventing head and brain injury amongst riders who crash. A case–control study conducted by Thompson et al. (1989) in Seattle over a period of 1 year found that bicycle helmets reduced the risk of head and brain injury by 85% and 88%, respectively. In a second larger case–control study by the same group (Thompson et al., 1996), helmets decreased the risk of head injury by 69%, brain injury by 65%, and severe brain injury by 74%. Helmets were found to be equally effective in accidents involving motor vehicles and those not involving motor vehicles. Furthermore, helmets were found to provide substantial protection from head injuries across all age groups. Amoros et al. (2012) recently conducted a case–control study in France and studied helmet effectiveness over more than 13,500 cyclist injuries. They concluded that helmets were associated with a decreased risk of head injury in cyclist trauma and this decrease seemed to be more pronounced for severe head injuries. Maimaris et al. (1994) studied over a thousand patients that sustained cycling-related injuries who were treated at an emergency department in England. They concluded that helmets reduced the risk of head injury by a factor of more than three. Heng et al. (2006) found that helmet use significantly reduced the risk of head and facial injury in a 2006 study of cycling trauma in Singapore. Despite the protection provided by helmets, as demonstrated by the epidemiological studies above, the safety benefits offered by helmets are not universally accepted. Many cities, towns, states and provinces do not have helmet laws and many cyclists do not wear helmets (Page et al., 2012). Anti-helmet groups state that helmets are not effective and that, in some cases, due to the increased size of a helmeted head compared to a bare head or due to the compliance of the shell or presence of vent holes, helmets can cause “rotational” injuries such as diffuse axonal injury (DAI). In the lay press, some groups claim that helmets cause injuries by obstructing vision or blocking sound. Researchers have also published articles, critical of the many epidemiological studies (cited above) that show that helmets are highly effective at preventing head injuries, accusing them of bias and conflicts of interest (Curnow, 2006; Elvik, 2011). Curnow argued that bicycle helmets are not as effective as claimed because previous epidemiological studies have not considered rotational injury (Curnow, 2003). There is considerable debate on the merit and limitations of the epidemiological evidence (Curnow, 2006, 2003; Elvik, 2011; Hagel and Barry Pless, 2006). One limitation of the epidemiological approaches is that it infers helmet performance during the impact from evidence collected after the impact and thus the severity of the head impact under study is never known. It is not our purpose to debate the merit of the epidemiological literature. Here we aim to explore the extent to which the epidemiological evidence of helmet efficacy can be supported or contradicted by
a biomechanical study that allows study of helmet performance during the impact. Biomechanical investigations of helmet efficacy, and indeed helmet certification standards, simulate helmeted head impact by dropping helmeted headforms onto prescribed impact surfaces. In helmet certification standards, the primary metric to assess impact management efficacy is linear headform acceleration measured during a drop test; helmets are considered to have met the certification criteria if the helmeted headform acceleration is below a prescribed threshold. The threshold varies from standard to standard (Table 1), and is not directly correlated to established risk curves. The standards generally require that helmets be certified using a magnesium headform. The range of drop heights associated with these standards is from 1.5 m (EN1078) to 2.2 m (Snell B95A) (Table 1). In biomechanical investigations, linear and rotational head accelerations are measured during the impact and helmet efficacy is determined by comparing these accelerations, and other derived metrics such as the Head Injury Criterion (HIC), to injury risk functions. For example, Mertz et al. have established head injury probability curves, in terms of HIC and linear acceleration, for the Hybrid III headform which was originally developed for automotive crash testing (Mertz et al., 2003). Because injury tolerances exist for this headform, the Hybrid III is increasingly applied in biomechanical helmet and head impact studies (Beckwith et al., 2012; Kendall et al., 2012; Pang et al., 2011; Pellman et al., 2003; Scher, 2006; Scher et al., 2009; Viano and Halstead, 2012; Viano and Pellman, 2005). Overall, the biomechanical studies indicate that helmets significantly reduce head accelerations relative to unhelmeted impacts (Benz et al., 1993; Hodgson, 1990; Mattei et al., 2012; Scher, 2006) or to impacts with thin uncertified novelty helmets (DeMarco et al., 2010; Scher et al., 2009). Furthermore, because linear head acceleration is known to be monotonically correlated to concussion and skull fracture risk (Greenwald et al., 2008; Mertz et al., 2003; Pellman et al., 2003) they are therefore known to reduce the risk of sustaining head injury. The biomechanical comparison that best matches the epidemiological studies, and thus that would be best able to augment the debate in that field, is a comparison of helmeted and unhelmeted head impact under identical impact conditions. Unfortunately, these tests are difficult to perform because of limitations of the magnesium head forms that are mandated in bicycle helmet standards and that have thus most often been used in bicycle helmet impact tests. The magnesium head forms are at high risk of damage if they are tested with no helmet and they have not been validated to match the expected human response for bare head impacts. Thus there is no test series available to our knowledge that contrasts helmeted and unhelmeted impacts for contemporary bicycle helmets under direct matched impact. Hodgson contrasted early 1990s era helmets with an unhelmeted impact using a “small humanoid headform”, Benz et al. dropped an unhelmeted Hybrid II headform from a lower height than their helmeted impacts and Mattei et al. dropped human cadaver skulls with and without helmets from six and nine inch drop heights (Benz et al., 1993; Hodgson, 1990; Mattei et al., 2012). All of these studies demonstrated a dramatic decrease in head accelerations for the helmeted compared to the
Table 1 Comparison of several bicycle helmet standards. Standard
Reference
Drop height (m)
Drop height (feet)
Criteria (g’s)
Consumer Product Safety Commission (CPSC) Snell Memorial Foundation (Snell) American Society for Testing and Materials (ASTM) Canadian Standards Association (CSA) European Standards (CEN)
16 CFR Part 1203 BF95 (1998 Revision) ASTM F1447F12 CSA D113 2FM89 (Reaffirmed 2004) EN 1078
2 2.2 2 1.6 1.5
6.6 7.2 6.6 5.2 4.9
300 300 300 250 250
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Fig. 1. Photograph showing helmeted Hybrid III headform (left) and unhelmeted Hybrid III headform (right) in contact with the steel anvil. The Hybrid III headform was attached to the ball arm which was mounted to a linear bearing on the monorail drop-tower.
unhelmeted impacts. However, to the best of our knowledge, there have been no studies that compare the crucial situation of helmeted and unhelmeted impacts using contemporary bicycle helmets, with a head form that is validated for bare head impacts and from drop heights that compare to bicycle helmet standards and real world cycling falls. The data that would result from such a test series is directly relevant and indeed central to the ongoing debate of bicycle helmet efficacy. The objective of this study was to assess the biomechanical efficacy of bicycle helmets to reduce risk of head injury in simulated head impacts from drop heights consistent with bicycle helmet standards and real world cycling head impacts. We designed and fabricated a custom-made test fixture that allowed us to attach a Hybrid III headform to a monorail drop tower and performed both helmeted and unhelmeted drops. The Hybrid III head form can be tested without a helmet and it is validated in bare head impacts (Foster et al., 1977). Linear head acceleration was measured, and HIC and injury risk were determined from these accelerations to ascertain the efficacy of helmets to reduce risk of head injury. 2. Methods and materials We simulated head impacts using a monorail drop tower similar to those specified in helmet certification standards. We fabricated a custom-made test fixture that allowed us to attach a Hybrid III headform (Humanetics Inc., Plymouth, MI, USA), that corresponded to a 50th percentile male head, to the drop tower. A ball-arm was mounted to a monorail drop tower that was purpose-built for this application. Paired tests were performed in order to study the risk of injury in helmeted and unhelmeted impacts. A paired test is defined as two drops onto an identical anvil and from identical drop heights both with and without a helmet. The impact surface for all drops was a flat, fixed steel anvil. Fig. 1 shows the anvil, helmeted and unhelmeted Hybrid III headform and features of the bearing and guide rail. Translational acceleration along the direction of impact was measured using a single axis accelerometer (±2000 g range, Endevco model 7264C2000, Meggitt Sensing Systems, San Juan Capistrano, CA, USA), which was mounted to the center of the ball-arm which in turn was placed within the Hybrid III head close to the head center of mass. The mass of the entire drop assembly, including the Hybrid III headform, ball-arm, and linear bearing was 5.05 kg. The mass of the helmet was approximately 0.25 kg and was considered additional mass for the helmeted drops. The helmeted and unhelmeted drops were conducted from nominal heights starting at 0.5 m to 3 m in 0.5 m increments. This range brackets heights used in certification standards but exceeds
the maximum height of typical standards (CSA D113.2-M89 (1.7 m), CPSC (2.0 m), ASTM F1447 (2.0 m), EN1078 (1.5 m) Snell B95A (2.2 m)) to allow study of higher energy impacts that can occur in real-world cycling where falls can happen while traveling at considerable speed. Our testing range also brackets the range of perpendicular impacts documented for reconstructed bicycle falls (Fahlstedt et al., 2012). Two drops were performed at 0.5 m, 1 m, 1.5 m, 2.5 m and 3 m; one drop for a helmeted Hybrid III headform and one unhelmeted. Six drops, three helmeted and three unhelmeted, were performed from 2 m. More drops were performed at 2 m than other heights to obtain drop speed and acceleration data that would allow a limited investigation of the repeatability of the experiment and to do so at common drop height used in bicycle helmet standards. To assess repeatability, we calculated the maximum difference in both peak acceleration and HIC and expressed these differences as a percentage of the mean peak acceleration and HIC. In total sixteen drops were conducted (8 helmeted, 8 unhelmeted). The headform was adjusted so that impacts took place to the forehead of the headform as seen in Fig. 1. Actual drop heights were increased by approximately 5 cm above the nominal drop height to account for friction in the drop rail. Speed at impact was calculated using high-speed video and was found to be within 5% of the expected velocity for each respective drop height. All helmets used in this work were CCM V15 Backtrail bicycle helmets (Reebok-CCM Hockey, Montreal, QC, Canada). The helmets were constructed with a micro-shell and an expanded polystyrene liner. The helmets conformed to the standards set out by the Consumer Product Safety Commission (CPSC) (CPSC, 1998). In impacts where helmets were used, the helmet was placed on the Hybrid III headform in a standardized fashion. The orientation of the headform was held constant for all drop heights using angle and position landmarks drawn on the Hybrid III headform (Fig. 1). The chin retention strap was tightened to secure the helmet to the Hybrid III headform (Fig. 1). The helmets were also equipped with a ratcheting tension system that is designed to pass inferior to the occipital protuberance. This was tightened prior to all drop tests. A check for helmet fit and secure attachment to the Hybrid III headform involved manipulation of the helmet on the head to ensure no visible relative motion. Each helmet was used for a single drop and then replaced with a new helmet. An Analog Devices (Analog Devices Inc., Norwood, MA) data acquisition system was used to collect the data, with the acceleration signal sampled at 39 kHz and hardware anti-alias filtered to comply with SAE J211-1 (“SAE J211 Instrumentation for Impact Test – Part 1: Electronic Instrumentation”). In addition, the accelerometer data were low-pass filtered at 1650 Hz (CFC1000) during post-processing as per SAE J211-1.
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Fig. 2. Typical acceleration data plotted versus time for both a helmeted and unhelmeted Hybrid III. Data shown is for the 2 m drop height. Acceleration is expressed in g, where one g corresponds to 9.81 m/s2 .
Peak head accelerations were used as a biomechanical metric of helmet efficacy and were compared for the paired helmeted and unhelmeted tests. In order to assess the risk of injury associated with these impacts, head accelerations were also compared to the Injury Assessment Reference Value (IARV), and probability curves published by Mertz et al. (2003). For presenting peak acceleration data, we use a 5% risk threshold for skull fracture based on peak acceleration (180 g). The Head Injury Criterion (hereafter HIC) was calculated during post-processing using Eq. (1). The HIC quantifies head impact severity by incorporating time of acceleration exposure and acceleration magnitude.
HIC15 =
1 t2 − t1
2.5
t2
a(t)dt t1
(t2 − t1 )
(1) max
For this analysis, a(t) is the head acceleration, in g, as measured by the single axis accelerometer, and the time interval (t2 − t1 ) was chosen to maximize HIC over a maximum duration of 15 ms (Eppinger et al., 1999). In the subject testing the sensing axis of the accelerometer was aligned with the direction of impact and thus captured the resultant acceleration. Similar to the acceleration analysis, HIC15 values were compared to the IARV of 700 which Mertz et al. have reported corresponds to a 5% risk of AIS ≥ 4 brain injury for the adult population (Mertz et al., 2003).
Fig. 3. Peak accelerations for both helmeted and unhelmeted drops. Numbers over bars indicate peak acceleration. For 2 m drop height, results stated are the mean value calculated from three drops. Horizontal dashed line indicates the IARVof 180 g (5% chance of skull fracture) for a midsize male.
probability were all of smaller magnitude in drops where the Hybrid III was helmeted. The duration of the impact pulse was larger in helmeted drops relative to unhelmeted. As the head decelerated, a small amount (approximately less than 5 mm) of sliding outward occurred between the helmet shell and the impact surface as the helmet shell and liner deformed during impact. Peak accelerations (Fig. 3) were smaller in helmeted drops relative to unhelmeted drops, for all drop heights. On average and considering all drops, the peak accelerations for helmeted drops were smaller by a factor of 4.2 relative to unhelmeted. For the severity of impacts tested, peak acceleration exhibited a linear relationship with drop height. In the unhelmeted situation, the head accelerations were above the IARV of 180 g for every drop from 0.5 m to 3 m. For drop heights of 0.5–1.5 m, helmets decreased the peak accelerations to a value below the IARV (Fig. 3). Fig. 4 shows maximum HIC for helmeted and unhelmeted drops from all heights. For each drop height, helmets reduced HIC relative to unhelmeted drops. The mean interval required to maximize HIC for unhelmeted and helmeted drops was 1.0 ms and 5.0 ms, respectively. The increased HIC interval for helmeted drops is consistent with the considerably wider (in the time domain) acceleration peak for helmeted drops relative to unhelmeted shown in Fig. 2. Fig. 5 shows the calculated risk of a severe brain injury (AIS 4+) for all helmeted and unhelmeted drops. Overall, the helmeted drops dramatically reduced the risk across all drop heights. For drops of 1 meter and greater, the unhelmeted condition resulted in essentially
3. Results Repeatability was evaluated by analyzing multiple drops at 2 m (3 helmeted drops and 3 unhelmeted drops). The maximum interdrop difference in peak acceleration was 1.5% and 3.3% (percentage of mean peak accelerations for 2 m drop) for the helmeted and unhelmeted Hybrid III headform, respectively. Similarly, maximum inter-drop differences in HIC were 6.0% and 5.0% (percentage of mean HIC), respectively. Fig. 2 shows typical acceleration curves plotted over the time of the impact event for both the unhelmeted and helmeted Hybrid III headform. In general, the acceleration magnitudes plotted over time exhibited a single abrupt increase in acceleration, which continues to the peak acceleration, followed by an abrupt decrease in acceleration. Following this, the accelerations fluctuate (e.g. Fig. 2, after 5 ms for the unhelmeted data) and these fluctuations correspond to head/helmet–anvil interactions that are secondary to the initial head/helmet-to-anvil impact (i.e. the head “bounces” off of the anvil). In general, peak accelerations, HIC and injury
Fig. 4. Head Injury Criterion (HIC) calculated using HIC15 convention for both helmeted and unhelmeted drops. Numeric values over bars indicate HIC values and long dashed line indicates IARV based on HIC.
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Fig. 5. Calculated probability of sustaining a severe brain injury (i.e. AIS 4+) based on the HIC values for the helmeted and unhelmeted drops. Numeric values over bars indicate risk magnitude.
100% chance of an AIS 4+ brain injury. For all drops of 2 m or less, the helmeted condition resulted in a risk of below 35%. However, at the 3 m drop height the risk of an AIS 4+ brain injury exceeded 90% for the helmeted condition. For each of the helmeted drops, various levels of helmet deformation and helmet shell and foam damage were evident. The expanded polystyrene (EPS) foam cracked near the impact location in all but the lowest drop height (0.5 m) and the micro-shell fractured in the 3 m drop. Upon further inspection, all of the test scenarios resulted in plastic deformation of the EPS foam liner near the impact location. 4. Discussion The overarching objective of this study was to assess the efficacy of certified contemporary bicycle helmets to mitigate skull and brain injury risk in head impacts with characteristic velocities matching impact velocities from helmet certification standards and also matching those from the biomechanical literature. We used a purpose-built drop tower with a Hybrid III test headform and a uniaxial accelerometer aligned to the direction of impact to measure linear head accelerations during both helmeted and unhelmeted drops. Linear head acceleration is one accepted mechanical measure that can be related to both skull injury risk and brain injury risk (Mertz et al., 2003; Pellman et al., 2003) both directly and through HIC (Mertz et al., 2003), and therefore acceleration was our primary biomechanical measure to assess efficacy. We have characterized the ability of one typical contemporary bicycle helmet to reduce the severity of a head impact and reduce the risk of severe life-threatening skull and brain injury, compared to not wearing a helmet, in matched impact tests where impact severities (i.e. drop height and pre-impact head velocity) were identical for the case of the helmeted and unhelmeted headform. The tested helmet dramatically decreased peak linear head acceleration (Fig. 3), HIC15 (Fig. 4) and the potential for severe brain injury (Fig. 5) in all impacts. Considering peak acceleration (Fig. 3), the helmeted headform experienced accelerations below the IARV of 180 g (Mertz et al., 2003) in drops from 0.5 m, 1.0 m, and 1.5 m while the unhelmeted headform experienced acceleration well above the IARV in drops from all heights. In drops from 2.0 m up to 3.0 m, the helmeted headform experienced accelerations above the IARV, but helmeted headform accelerations were at least 4 times smaller than those of the unhelmeted headform (Fig. 3). Evaluation of HIC15 (Fig. 4) was consistent with the acceleration results when compared with the IARV of 700. Probability of skull fracture (not shown)
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and severe brain injury (Fig. 5) is reduced, for all drop heights, for the helmeted headform relative to the unhelmeted headform. This biomechanical evidence clearly indicates that contemporary bike helmets are highly effective at reducing injury risk through paired helmeted and unhelmeted impacts with realistic drop heights and impact speeds. For example, the helmets reduced the head peak acceleration from 824 g to 181 g for drops of 2.0 m reducing the risk of skull fracture from 99.9%+ to 5%. The reductions in head acceleration and HIC15 described above demonstrate that certified helmets significantly reduce the risk of sustaining severe and even fatal injuries. It is worth noting that common helmet standards do not seem to be designed to prevent head accelerations from exceeding the IARV values for severe skull and brain injuries published by Mertz et al. (2003), although it is noted that the helmets reduced the accelerations to a value well below that specified in the standard. The helmeted drop from 2.0 m resulted in a HIC of 1250 that corresponds to a 34% chance of severe brain injury. It may be necessary to re-evaluate the helmet standards and contemplate lowering the allowable accelerations in some standards and to require testing from multiple drop heights in order to decrease the potential for serious and severe skull and brain injuries from falls of various drop heights. Common injuries coded AIS 4 (severe) and above include penetrating skull injuries leading to brain injury, large contusions (e.g., coupe-contrecoup, intermediate, and gliding) and hematomas (e.g., subdural, subarachnoid, and intracerebral), as well as diffuse axonal injury with associated loss of consciousness for a period exceeding six hours. The severity of these injuries when coded as 4+ range from severe to maximum (usually fatal) (Gennarelli and Wodzin, 2005). Considering a realistic bicycle accident scenario documented in the literature (Fahlstedt et al., 2012) where a cyclist was thrown at 20 km/h (i.e. 5.6 m/s which corresponds to a drop height of approximately 1.5 m), our analysis indicates that a helmeted cyclist in this situation would have a 9% chance of sustaining the severe brain and skull injuries noted above whereas an unhelmeted cyclist would have sustained these injuries with 99.9% certainty. In other words, a helmet would have reduced the probability of skull fracture or life threatening brain injury from very likely to highly unlikely. Evaluation of the 3 m drops demonstrate that helmets only offer a finite amount of protection. However, the 7.7 m/s impact speed is not representative of most real-world bicycle impacts (Fahlstedt et al., 2012). At impact speeds of this velocity the energy management capability of the helmet is saturated (Newman, 2002) and the EPS liner bottoms out. The range of drop heights that we tested was consistent with the range of impact speeds that has been documented as plausible for cyclist impact scenarios. Our 3.0 m, 2.5 m, and 2.0 m drops were consistent with the resultant cyclist head velocity in the Fahlstedt study that was oblique to the ground (i.e. glancing off it). However, if the cyclist hit a curb or the wheel of a car at the point of glancing off the ground then there would be head impact of similar severity and impact direction to the higher impact severity tests that we carried out. The biomechanical results for acceleration and HIC15 , and the corresponding results for reduced risk of severe skull and brain injury, complement epidemiological studies that have sought to assess the protective efficacy of bicycle helmets. The majority of the epidemiological studies suggest that helmets are effective at reducing injury risk in a range of sports and across both adult and youth segments of the population. However, concerns have been raised in recent years over issues of selection-bias, failure to compensate for time-trending and public policy, and lack of control for confounding aspects of the input data for statistical studies including type and mechanism of head/brain injury and severity of the head impact (Curnow, 2006; Elvik, 2011). A fundamental reason for this study is that the lack of appropriate knowledge of and
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statistical control for the severity of the impact confounds the results and decreases the significance of results below what they would be if appropriate controls were applied. As a consequence, the results in some studies suggest helmets are only slightly effective at preventing head injuries or, in extreme cases, actually cause head injuries (Elvik, 2011). However, the results of our biomechanical study, where impact severity was controlled in helmeted and unhelmeted impacts, strongly refute the epidemiological studies that suggest helmets are marginally effective. Indeed, in all impacts the risk of sustaining injury was reduced when the Hybrid III headform was helmeted, and in the case of a realistic bicycle head impact (5.5 m/s impact speed which corresponds to a drop height of approximately 1.5 m), a helmeted cyclist would have a 9% chance of sustaining AIS 4+ injuries whereas an unhelmeted cyclist would almost certainly sustain these injuries. The results of this study are in good agreement with previous biomechanical research on the protective efficacy of bicycle helmets. For example, Benz et al. (1993) conducted a study of child and adolescent helmets in drops from 1 m to 1.5 m using an unspecified “child headform” (Benz et al., 1993) and a Hybrid II headform for adolescent helmets, both dropped onto a flat anvil, and reported an overall threefold decrease in HIC as a result of protecting the headform with a helmet. Hodgson (1990) conducted a study using a Hodgson-WSU headform dropped from 1 m and 2 m onto flat and convex surfaces (Hodgson, 1990). For 1 m drops, the unhelmeted headform was at increased risk of injury (70–99% of the population would be injured) relative to the helmeted headform (
