A Comparison of Peak Linear and Angular Headform Accelerations Using Ice Hockey Helmets

Transcription

1 Journal of ASTM International, Vol. 6, No. 1 Paper ID JAI Available online at P. Rousseau, 1 A. Post, 1 and T. B. Hoshizaki 1 A Comparison of Peak Linear and Angular Headform Accelerations Using Ice Hockey Helmets ABSTRACT: The objective of this study was to quantify the ability of hockey helmets to manage peak angular and linear acceleration of the center of gravity of a Hybrid III headform at six impact locations and three velocities 5, 7, and 9 m/s. The method was intended to represent a reconstruction of helmet, shoulder, and elbow strikes to the head. Six different ice hockey helmets were impacted using a pneumatic linear impactor at velocities similar to those encountered in ice hockey. The results were then compared to impacts to an unhelmeted headform. The data demonstrated that under certain conditions, peak angular accelerations were above estimated injury thresholds, while peak linear accelerations remained below. This confirmed the importance of measuring peak angular acceleration and peak linear acceleration in order to better predict minor traumatic brain injuries. KEYWORDS: angular acceleration, ice hockey, helmet, mtbi Introduction Athletes performing in the National Hockey League NHL are bigger and stronger than 20 years ago 1. Combined with high skating velocities km/h 2, risks of sustaining an injury are high for ice hockey players. Reports have shown an increase in minor traumatic brain injuries mtbi in the NHL over the last 10 years 3. This type of injury represents 18 % of all reported ice hockey injuries 4 and is considered a dangerous injury, often requiring prolonged periods of time away from the game 5. Despite the introduction of helmets, originally designed to prevent severe head injuries, the incidence of mtbi has remained relatively stable for the past 30 years 5 7. Additional equipment, including visors and mouth guards, were believed to mitigate the injury; however, further research reported inconsistent effects on the incidence of mtbi Thus, helmets remain the most important piece of equipment in reducing the incidence of mtbi in ice hockey. Helmet performance is primarily influenced by safety standards which establish testing thresholds used to prevent major traumatic brain injuries including skull fractures and intracranial bleeds. Peak linear acceleration is often the only brain injury indicator used in these standards despite many studies demonstrating the importance of angular acceleration in predicting mtbi The test methodology most commonly used requires dropping a head form from a predetermined height on a fixed surface, simulating the mechanics of falling to the ice or into boards. This mechanism offers little compliance and does not account for impacts between two players, which are responsible for % of mtbi 5,6,16. Moreover, impacts away from the center of gravity, i.e., eccentric impacts, are not assessed. The objectives of this study were to quantify the ability of hockey helmets to manage peak angular acceleration and peak linear acceleration of the center of gravity of a Hybrid III headform at six impact locations and three velocities 5, 7, and 9 m/s. This will provide important insight into impact mechanics, crucial for head trauma prevention. Methodology The Pneumatic Linear Impactor System The pneumatic linear impactor system Fig. 1 consisted of three major components: 1 the support/piston frame, 2 the impacting arm, and 3 the table housing the Hybrid III headform. The frame supported the Manuscript received May 7, 2008; accepted for publication November 14, 2008; published online December Neurotrauma Impact Laboratory, University of Ottawa, School of Human Kinetics, Ottawa, Ontario, Canada Copyright 2009 by ASTM International, 100 Barr Harbor Drive, PO Box C700, West Conshohocken, PA

2 2 JOURNAL OF ASTM INTERNATIONAL FIG. 1 Linear impactor. impacting arm, the compressed air canister, and the piston which was controlled electronically. The impacting arm mass kg was propelled horizontally by compressed air. On the tip of the impacting arm was a cap consisting of a kg hemispherical nylon pad with a mm thick vinyl nitrile 602 foam disk underneath. The mobile table kg was installed at the receiving end to support a Hybrid III head kg and neck kg and had a total mass of kg Fig. 2. The table allowed the dummy to slide backwards, thus allowing it to react in a realistic manner after the impact. A spring loaded brake system provided a safe stop following a displacement of m. The base supporting the Hybrid III head and attaching it to the moving table was built to allow for complete control over the location of impact. It could be adjusted in five degrees of freedom, including fore-aft x, lateral y, and up-down z translation, as well as fore-aft y and axial z rotation of the neck base. The adjustments were lockable and remained fixed throughout the testing. The Hybrid III Headform A Hybrid III 50th percentile male headform Fig. 3 was used in this study. The sensors mounted inside the Hybrid III headform were nine single-axis Endevco 7264C-2KTZ accelerometers, measurement range 500 peak g. They were positioned in an orthogonal arrangement following the array 17. FIG. 2 Mobile table supporting the Hybrid III headform.

3 ROUSSEAU ET AL. ON LINEAR AND ANGULAR HEADFORM ACCELERATIONS 3 FIG. 3 Hybrid III 50th percentile male headform (mass kg). The processing of the nine signals allowed the determination of the complete three-dimensional motion of the center of gravity of the head. The accelerations were collected at a frequency of 20 khz. The Helmets Six different hockey helmets representing three of the leading manufacturers were tested Fig. 4. The average mass for each model can be found in Table 1. The protective foams used in the helmets were either vinyl nitrile or expanded polypropylene. Test Procedure Three different hockey helmet models using vinyl nitrile foam and three models using expanded polypropylene were impacted using a pneumatic linear impactor. The helmets were struck three times per impact location using a virgin helmet at each of the following three velocities 5, 7, and 9 m/s. The unhelmeted headform, however, was not impacted at 9 m/s to prevent damaging the testing equipment. The average time between impacts was min. The impact locations were the following: 1 Front, center of gravity: The headform was directly facing the linear impactor 0 rotation. The impacts were located 30 1 mm above the intersection of the longitudinal plane and the reference plane Fig Front, 2.5 in. lateral translation: The headform was directly facing the linear impactor 0 rotation and was translated 2.5 in. to the right. The impacts were located 30 1 mm above the reference plane.

4 4 JOURNAL OF ASTM INTERNATIONAL FIG. 4 Standard ice hockey helmets; frontal view of a standard helmet (top left), interior view of a standard helmet (top right), frontal view of a standard helmet (bottom left), interior view of a standard helmet (bottom right). 3 Front, 5 in. lateral translation: The headform was directly facing the linear impactor 0 rotation and was translated 5 in. to the right. The impacts were located 30 1 mm above the reference plane. 4 Front Boss, center of gravity: The headform was rotated 52.5 towards the right and was translated 1.25 in. to the left to ensure an impact through the center of gravity. The impacts were located 30 1 mm above the reference plane. 5 Front Boss, 1.25 in. lateral translation: The headform was rotated 52.5 towards the right and was translated 2.5 in. to the left 1.25 in. offset. The impacts were located 30 1 mm above the reference plane. 6 Front Boss, 2.5 in. lateral translation: The headform was rotated 52.5 towards the right and was translated 3.75 in. to the left 2.5 in. offset. The impacts were located 30 1 mm above the reference plane. Following each impact, peak linear accelerations and angular were computed using a TDAS Pro Lab system DTS, Seal Beach, CA. A SAE J211 class 1000 filter was used on the data obtained from the accelerometers. All further data analysis was conducted by Bioproc 2 developed by Dr. Robertson, University of Ottawa. All peaks were compared statistically using ANOVAs. The impact velocities of the linear impactor were measured using a time gate and recorded by computer using National Instruments VI-Logger. The time gate was validated using a High Speed Imaging PCI-512 Fastcam. The Fastcam recorded the impact at a frequency of 2000 Hz using Photron Motion Tools. The video of the impactor arm striker was then digitized prior to impact to produce distance traveled over time which was used to calculate impact velocity. TABLE 1 Helmet weights. Model Foam Weight kg I II III

5 ROUSSEAU ET AL. ON LINEAR AND ANGULAR HEADFORM ACCELERATIONS 5 FIG. 5 Reference planes of a DOT FMVSS 218 headform [18]. Brain Injury Probability The results were compared to brain injury thresholds proposed by Zhang and colleagues 15 to evaluate the helmets capability to manage peak accelerations. These thresholds were obtained using reconstructions of helmet-to-helmet field collisions recorded during National Football League games. Peak linear accelerations were reported to be 66, 82, and 106 G for a 25, 50, and 80 % probability of sustaining an mtbi, respectively. Peak angular accelerations were reported to be 4600, 5900, and 7900 rad/s 2 for a 25, 50, and 80 % probability of sustaining an mtbi, respectively 15. Results Tables 2 4 present the average 1 standard deviation peak linear and angular accelerations for front impacts at 5, 7, and 9 m/s. Results demonstrated that, at all velocities, impacts through the center of gravity and with a 2.5 in. translation generated higher peak linear and angular accelerations then impacts with a 5 in. translation p Impacts through the center of gravity produced higher peak linear accelerations than impacts with a 2.5 in. translation for impact at 5, and 7 m/s p Impacts through the center of gravity and impacts with a 2.5 in. translation produced similar peak angular accelerations with the exception of impacts at 7 m/s p TABLE 2 Average (standard deviation) peak linear acceleration and peak angular acceleration for front impacts with 0 (CofG), 2.5, and 5 in. lateral translation at 5m/s. Acceleration No helmet Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular 11, rad/s in. Linear g Angular rad/s in. Linear g Angular rad/s

6 6 JOURNAL OF ASTM INTERNATIONAL TABLE 3 Average (standard deviation) peak linear acceleration and peak angular acceleration for front impacts with 0 (CofG), 2.5 and 5 in. lateral translation at 7m/s. Acceleration No helmet Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular 14, rad/s in. Linear g Angular rad/s in. Linear g Angular rad/s Tables 5 7 show the average 1 standard deviation peak linear accelerations and peak angular accelerations for front boss impacts at 5, 7, and 9 m/s. Results show that impacts through the center of gravity generated lower peak linear accelerations then impacts with a 2.5 in. translation at 9 m/s p They also produced lower peak accelerations then impacts with a 1.25 in. translation at 7 and 9m/s p Impacts through the center of gravity generated lower peak angular accelerations then impacts with a 1.25 and 2.5 in. translation at all velocities p Impacts with a 1.25 in. translation and impacts with a 2.5 in. translation produced similar peak linear and angular accelerations at all velocities. TABLE 4 Average (standard deviation) peak linear acceleration and peak angular acceleration for front impacts with 0 in. (CofG), 2.5 and 5 in. lateral translation at 9m/s. Acceleration Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular rad/s in. Linear g Angular rad/s , in. Linear g Angular rad/s TABLE 5 Average (standard deviation) peak linear acceleration and peak angular acceleration for front boss impacts with 0 (CofG), 1.25 and 2.5 in. lateral translation at 5m/s. Acceleration No helmet Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular 12, rad/s 2 1, in. Linear g Angular 10, rad/s in. Linear g Angular rad/s

7 ROUSSEAU ET AL. ON LINEAR AND ANGULAR HEADFORM ACCELERATIONS 7 TABLE 6 Average (standard deviation) peak linear acceleration and peak angular acceleration for front boss impacts with 0 (CofG), 1.25 and 2.5 in. lateral translation at 7m/s. Acceleration No helmet Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular 18, rad/s in. Linear g Angular 14, , rad/s in. Linear g Angular ,411 rad/s Figure 6 shows peak linear and angular accelerations for front impacts through the center of gravity at 5, 7, and 9 m/s. At 5 m/s, peak linear and angular accelerations were below a 25 % and above an 80 % probability of sustaining mtbi, for helmeted and unhelmeted headform impacts, respectively. At 7 m/ s, peak linear accelerations for two helmets were above the 50 % probability of injury. One of the above helmets was also above the 50 % injury risk for angular acceleration; three were above the 25 % threshold, while two helmets were below. Peak linear and angular accelerations were above an 80 % probability of sustaining an mtbi. At 9 m/s, all peak linear accelerations were above the 80 % probability of sustaining mtbi. Peak angular accelerations showed more variation as two helmets were above 80 % risk of injury, two were above 50 %, and two were above 25 %. Figure 7 reveals peak linear and angular accelerations for front boss impacts with a 2.5 in. translation at 5, 7, and 9 m/s. At 5 m/s, all peak linear accelerations were below the 25 % probability of sustaining an mtbi for helmeted impacts and below 50 % for unhelmeted impacts. For peak angular accelerations, three helmets were above the 25 % risk of injury, two were above the 50 %, while the unhelmeted head was just below 50 %. At 7 m/s, the unhelmeted head as well as one impact with a helmet had a peak linear acceleration above the 50 % probability of injury, while the others remained below. At the same velocity, four helmets were above the 80 % injury risk for angular acceleration, while the unhelmeted head and the remaining two helmets were below. At 9 m/s, all peak accelerations were above the 80 % probability of sustaining an mtbi. Discussion The results demonstrated the relevance of measuring peak linear and angular acceleration when monitoring mtbi risk. As reflected in Tables 2 7, helmets certified using the same tests did not perform similarly. In some conditions, helmets with similar peak linear accelerations had different peak angular accelerations. TABLE 7 Average (standard deviation) peak linear acceleration and peak angular acceleration for front boss impacts with 0 (CofG), 1.25 and 2.5 in. lateral translation at 9m/s. Acceleration Helmet A Helmet B Helmet C Helmet D Helmet E Helmet F CofG Linear g Angular ,217 rad/s in. Linear g Angular 13,350 12,936 11,776 14,058 16,390 14,488 rad/s , in. Linear g Angular 11,042 12,366 13,660 16,456 16,483 16,758 rad/s ,

8 8 JOURNAL OF ASTM INTERNATIONAL FIG. 6 Peak linear acceleration and peak angular acceleration for six ice hockey helmets as well as an unhelmeted head during a direct impact to the front of the headform, through the center of gravity. The three lines represent a 25, 50, and 80 % probability of sustaining a mtbi. FIG. 7 Peak linear acceleration and peak angular acceleration for six ice hockey helmets as well as an unhelmeted head during an impact to the front boss of the headform (52.5 rotation), with a 2.5 in. translation to the left. The three lines represent a 25, 50, and 80 % probability of sustaining a mtbi.

9 ROUSSEAU ET AL. ON LINEAR AND ANGULAR HEADFORM ACCELERATIONS 9 Furthermore, a review of Table 2 indicates a large difference in the angular acceleration response of helmets C and D. This variation may be attributed to differences in shell geometry, foam type, and foam density. Further research is required to understand the effects of helmet characteristics on peak linear and angular accelerations. The data also showed that helmets capable of maintaining peak linear accelerations below an acceptable risk of injury in certain conditions were incapable of doing so for peak angular accelerations Fig. 7. Peak angular accelerations reached during front impacts through the center of gravity remained below the 50 % risk of injury Fig. 6. Conversely, peak angular accelerations reached during front boss impacts with a 2.5 in. translation were above the 80 % risk of injury Fig. 7. Thus, impact location had an influence on the magnitude of peak angular acceleration, which is in accordance with similar studies When the results were compared to the ones obtained using the headform alone, it became apparent that the ability of the helmets to decrease peak linear and angular accelerations was influenced by the impact characteristics. An analysis of front boss impacts with a 2.5 in. lateral translation impact Fig. 7 showed that all six helmets were ineffective at reducing peak angular acceleration. The higher values suggest that present helmets cannot protect the head in such situations and may even put the athlete at a greater risk of sustaining an mtbi. These results highlight the importance of adding eccentric impacts to current test protocols. Factors such as shell size and geometry should be considered when designing helmets as they most certainly have an effect on resulting angular acceleration. The results also illustrate the importance of including peak angular acceleration when evaluating the performance of ice hockey helmets. It was shown that peak angular acceleration did not always react in the same manner as linear acceleration and that it identified certain impacts as being likely to provoke higher angular accelerations, even when peak linear accelerations remained low. To our knowledge, no other study has compared peak angular and linear accelerations in ice hockey helmets. Nonetheless, studies on American football helmets conducted by Pellman and his colleagues, reported similar trends There were a few limitations to this study. A 50th percentile male Hybrid III head and neck were used as human surrogates. Although the Hybrid III is the most widely used mannequin, it is not biofidelic, meaning that it does not imitate the human head s exact dynamic properties 25,26. Furthermore, the neck was only calibrated for rotation around the y axis. Nevertheless, the Hybrid III dummy remains the industry standard and has a multiaxis neck capable of withstanding the forces generated by the highvelocity impacts used in this study. It is important to keep in mind that these results were achieved using a pneumatic linear impactor with an effective striking mass of kg, which can explain why impacts at 7 and 9 m/s showed such high risks of injury. It is understood that effective mass plays a large role in impact mechanics; however, this study was not meant to be an exact reconstruction of impacts experienced in ice hockey. Finally, the effect of the sliding table supporting the headform is on the dynamic response of the head form is not known. For this study, the table offered very little resistance to motion as it allowed the dummy to slide backwards following the impact. Its effects on impact mechanics have not yet been fully studied. Conclusion The ice hockey helmets tested in this study showed a limited ability to manage peak angular accelerations and peak linear accelerations below defined thresholds under all six conditions. Furthermore, peak linear and angular accelerations did not respond uniformly across impact location and across impact velocities. Thus the following recommendations are made: 1 Peak angular accelerations should be considered when designing and evaluating the performance of ice hockey helmets. 2 Impacts locations away from the center of gravity should be added to existing test protocols. Acknowledgments The authors would like to thank Itech-Mission, Nike-Bauer, and Reebok-CCM for kindly providing them with helmets. They would also like to thank Xenith for supporting their laboratory.

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