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David C. Viano, Ph.D., Joseph D. McCleary, Dennis V. Andrzejak, and David H.
Janda, M.D.
Institute for Preventative Sports Medicine and Orthopedic
Surgery Associates, Ann Arbor, Michigan, U.S.A.
See the Abstract
It has been estimated that > 19 million children are involved in youth
baseball in the United States. It has also been estimated that softball and
baseball cause more injuries leading to emergency room visits in the United
States than any other sport. Between 1983-9, > 2.6 million injuries were
documented through selected emergency rooms throughout the United States by the
Consumer Product Safety Commission secondary to baseball and softball. Although
this figure is an underestimate because it does not include nonhospitalization
physician visits, it does indicate the magnitude of the current problem. The
Consumer Product Safety Commission has also concluded that more baseball related
fatalities occur in the 5-14-year-old age group than in any other sport. The two
scenarios documented by the Consumer Product Safety Commission focus on the
impact of a baseball to a child's chest and the impact of a baseball to a
child's head. Significant improvements have been made in the safety of organized
sports, including the use of helmets by batters and base runners in baseball,
break away bases in baseball and softball, and protective eye wear in racquet
sports. With a greater interest in sport safety, many new products have been
developed and marketed to reduce potential injury risks. Unfortunately, there
are only a few certification tests that can be used to quantify the relative
safety provided by various products.
Recently, softer baseballs have been promoted as substantially reducing the
risks of head injury for unprotected, nonhelmeted baseball players. Studies have
been conducted comparing the risks of head injury between conventional hardballs
and the softer baseballs. The data have been widely distributed in promoting a
shift to softer baseballs in youth sport.
Today, helmets are routinely worn for head protection in organized baseball
games while batting and base running. This has substantially improved the level
of safety. Advances have recently been made in the construction of baseballs,
and softer balls are being marketed as a further safety measure. The balls have
been evaluated using the National Operating Committee on Standards for Athletic
Equipment (NOCSAE) helmet evaluation test. This procedure has only limited
validation for direct head impacts and shows a significant safety improvement
with the softer baseballs for an unprotected head impact. We used the
5th-percentile Hybrid III female dummy, which is similar in size to a
10-12-year-old child, for head impacts with various baseballs at 27 m/s (60
mi/h). These tests provide an independent evaluation of softer baseballs and a
check of the current NOCSAE standard as a safety test for unprotected head
impacts.
Our study is part of a general evaluation of the effectiveness of baseball
safety equipment using in dependent tests and advanced methods to objectively
quantify levels of protection with the equipment. The study draws on the
practical experiences of automotive safety scientists and clinicians, and
applies the most advanced biomechanical tools and surrogates to address the
protection of children playing baseball.
Our previous studies of baseball safety have primarily addressed fatalities
occurring from chest impact by a batted or pitched baseball. The epidemiology
and pathophysiology of chest impact deaths have been reported by Viano et al.
[33] based on an earlier unpublished report to the
Consumer Product Safety Commission. That study indicated that normal healthy
children were vulnerable as batters (primarily turning to bunt), pitchers, or
fielders struck directly on the torso. Another study by Viano et al. [35] developed a model of the chest impact death and
determined the mechanisms of fatal injury. Children may succumb to fatal
dysfunction of heart and respiratory rhythm due to the high speed of baseball
impact and shock to the cardiorespiratory systems. Also noteworthy was the
apparent failure of emergency cardiopulmonary resuscitation procedures in
reviving the children.
A final study in the series by Janda et al. [8]
addressed the qualification of surrogates to study potential approaches to
preventing chest impact deaths in baseball. By using the most advanced
automotive test dummy, the recently developed Hybrid 111 5th-percentile female,
and a specially modified 6-year-old child dummy, various pads, layers of padding
with hard shells, and softer baseballs were evaluated for their protective
function against life-threatening impact of the chest. The study found that most
approaches currently being considered to reduce the risk of chest injury provide
minimal protection and, in some cases, may aggravate the risk of injury.
On the other hand, various passive preventative measures have been instituted
within the sports medicine field and have yielded enormous reductions in injury
rates as well as an enormous reduction in health care costs. Peterson's analysis
of in jury rates in football led to the elimination of cross body blocking and
hence a substantial reduction in football injuries [24]. The institution of passive preventative measures
as illustrated by the research of Pashby et al. [23]
involving protective eye wear during racquet sports has led to a significant
diminution in eye injuries and associated health care costs.
Janda et al. [9,10]
have also studied the benefit of break-away bases in reducing the risk of ankle
in juries from sliding and base running. An independent analysis by the Centers
for Disease Control in regard to break-away base research in recreational
softball has led to a projected prevention of 1.7 mil lion injuries a year in
the United States, with a savings in health care costs of $2 billion/year nation
ally. These studies are illustrative of the beneficial effect passive
preventative measures can have both from a reduction in injury and health care
cost containment standpoint.
The earliest testing of helmet effectiveness in reducing the risks of head
injury involved metallic head forms. Acceleration of the head form in
standardized tests was used to qualify the safety of the helmets for civilians
(I) and military personnel [20]. While this approach
advanced the understanding of head protection, the lack of biofidelity
(similarity to the human impact response) was identified as an important
deficiency by Saczalski et al. [26], Slobodnik and
Nelson [28], and Slobodnik [27].
Contemporaneous work by Hodgson and co-workers [5,6] resulted in an improved
head form for helmet evaluations. Their studies demonstrated better biofidelity
and the procedure provided a refinement in evaluating helmet systems for head
impact protection. The Wayne State University (WSU) procedure was later adopted
by the NOCSAE [21] for baseball helmet testing.
While the work of Hodgson et al. at WSU has advanced safety in sports and
promoted better evaluation procedures for protective equipment, simultaneous
advances in the automotive industry have led to more widely accepted
anthropomorphic test dummies [16,18] and injury criteria [36] for the evaluation of passenger safety in motor
vehicle crashes. In fact, Melvin [15] conducted
baseball impact tests of the Part 572 and Hybrid III dummy in the 80-90 mi/h
speed range.
Worldwide advances in crash test surrogates include dummies with biofidelity
in head impact response, neck compliance, chest and abdomen deflection, and
extremity bone and joint injuries. The family of Hybrid III dummies includes the
50th percentile and 95th-percentile male, 5th-percentile female, and the yet to
be completed 6- and 3-year old child dummies. Each surrogate has impact
responses that have been scaled to the size of the dummy and appropriate body
anthropometry and inertial characteristics.
With the substantial technical advances in auto motive test dummy
biofidelity, there may be a logical transition of some civilian and military
impact standards to the use of Hybrid III dummies. This would also enable use of
further refinements in the family of test dummies, large data bases of injury
assessment reference values, and continuing improvements in the measurement of
three dimensional head translational and rotational acceleration [31,37], finite element and
other models of brain responses [11,12, 14,30], and neural trauma research [2, 3,13].
The latest revision of the NOCSAE helmet impact standard represents the
seventh modification of the test procedure and acceptance criterion [22]. However, the procedure continues with the head,
neck, and trolley test system originally developed by Hodgson et al. This study
explores the impact response of the 5th-percentile female Hybrid III dummy to
direct baseball impact on the front and side of the head. Responses from this
most advanced test device are compared to results using the NOCSAE protocol and
distributed to the public by Hodgson and Thomas [7]
and Heald [4].
METHODOLOGY
Hybrid III Dummy
The most advanced small test dummy available to
the safety community was utilized in this study. The 5th-percentile Hybrid III
female was selected be cause of its smaller size in comparison to the male test
dummies. The Hybrid III female weighs 110 lb (50 kg), with a head weight of 8.1
lb (3.68 kg). It has a head circumference of 53.3 cm, width of 14.5 cm, and
length of 18.3 cm [19]. These head dimensions are
comparable to those of a 10-year-old child [29],
whereas the overall body size is more representative of a 12-year-old child.
These dimensions make the device an excellent surrogate to test sports equipment
for child safety. The biofidelity of the Hybrid III dummy head has been
previously documented in the literature [17].
The 5th-percentile Hybrid III female dummy has human impact responses for
head impact, neck ex tension and bending, chest deflection, and many other body
characteristics and responses. The de vice is ideal for the study of baseball
impacts, al though there has been only a limited qualification of the Hybrid III
female for baseball loading of the chest at this time. Recent evaluations of the
frontal and side impact response of the Hybrid III head indicate acceptable
performance should be expected for direct impact by hard or compliant
surfaces.
While Hybrid III-type child dummies are being developed to represent 6- and
3-year-old children, they are not yet available for testing [18]. Once these devices are available, the 6-year-old
child model would make a suitable additional test device to evaluate baseball
impact protection. However, the 5th-percentile female model, with a head size
comparable to that of a 10-year-old child and body dimensions of a 12-year-old
child may ultimately be the most appropriate device since the median age of
baseball fatalities found by Viano et al. [33] was
between 10-11 years old.
Test setup
The female dummy was seated upright in front of a
pneumatic cannon that propelled the baseballs at a controlled speed (Fig. 1). The dummy was loosely supported so its whole-body
impact response would be representative of that of a freely standing child. This
approach uses the entire dummy and would be more realistic than the
head-neck-trolley system specified in the NOCSAE standard when the neck and
whole-body responses are an important factor in the head biomechanical response.
The pneumatic baseball gun was positioned 30 cm from the dummy's head and the
speed of impact was determined by a dual-laser-beam velocity system positioned
several centimeters in front of the baseball thrower. Upon baseball impact, the
head, neck, and whole body moved in response to the energy of the ball, which
rebounded off the head. High-speed videorecording was used to document the
squareness of the head impacts, the ball interaction with the head, and the
rebound.
The speed of baseball impact was 27 m/s (60 mi/h) as specified in the current
NOCSAE standard for helmet impact protection. Baseball impacts were delivered to
three head locations: frontally on the forehead, frontally between the eyes, and
laterally on the right temple above the level of the nasion. The precise
location of impact was related to the center of gravity (cg) of the head and the
placement of the seismic masses of the triaxial accelerometers in the head. The
temple impacts were within 6 mm of the head cg. This put the impact just above
the level of the nasion.
Fig. 1. Test setup for baseball impact of the 5th-percentile Hybrid III
female dummy. The pneumatic ball thrower is composed of a pressure tank, quick
release solenoid valve, and baseball guide tube. The dummy is positioned in
front of the ball thrower and loosely supported to simulate the head and whole
body responses from the impact (modified from reference 34).
Pneumatic baseball gun
Figure 1 also shows
the pressure-charged tank and quick-release solenoid valve system developed to
control the speed of baseball impact. The gun was used in our previous studies
and a relationship between tank pressure and speed of baseball impact has been
provided [34].
For each test, the gun barrel was loaded with an 80-mm-thick, 14-g foam wad
that tightly fit the tube and provided a uniform surface between the pressure
pulse and the spherical ball. The wad was compliant enough to conform to the
geometry of the ball but stiff enough to maintain a tight fit. The wad and ball
were placed into the barrel to the end connected to the solenoid valve. The tank
was charged with an air pressure of about 20 lb/sq. in (140 kPa) to achieve the
prescribed baseball speed. Tank pressure was adjusted slightly to maintain test
speed for the lighter baseballs. The fast-acting solenoid valve triggered the
impact.
Baseballs tested
Commercially available baseballs were tested in
the three-head impact configurations. Two tests were conducted at each location
and with each ball. Table 1 lists the baseballs,
manufacturer, weight, and relative kinetic energy using the official hard ball
weight (146.1 g) and speed (27 m/s) as standard. The relative impact energies
varied from 0.53 to 1.08. This range is identical to the relative weight of the
baseballs.
Table 1. Weights and energies of
baseballs tested
Kinetic
Baseball Weight, energy, Relative
type Company (g) (J) energy
Official Rawlings 146.1 52.58 1.000
RIF10 Worth 157.0 56.51 1.075
RIF5 Worth 154.4 55.57 1.057
RIF1 Worth 151.3 54.46 1.036
ProBall Kenko 146.8 52.84 1.005
SafeBall Western 104.8 37.72 0.717
SofTouch Easton 102.6 36.93 0.702
IncrediBall Easton 93.7 33.72 0.641
Orange Kenko 77.0 27.71 0.527
RIF, reduced injury factor.
(a) Kinetic energy assumes a speed of 26.83 m/s (60 mi/h)
for the various baseball impact.
(b) Relative energy was determined using a 146.1-g baseball
and a speed of 26.83 m/s (60 mi/h).
Biomechanical responses
Measurements were made of head impact
responses. Each axis of triaxial acceleration was re corded from the attachment
near the center of gravity of the dummy head. A triaxial neck load cell measured
forces at the junction between the head and neck. These responses were recorded
using a Dell 386 computer with an RC Electronics A-to-D conversion board. The
digitization rate was 10,000 samples/s/channel. The data were subsequently
filtered utilizing Society of Automotive Engineers Channel Class 1,000 Hz and
analyzed for the head injury criterion (HIC). The HIC algorithm optimizes a time
interval for an integral of head acceleration raised to the power of 2.5. The
impact speed of each test was also recorded.
Injury risk assessment
The peak resultant triaxial head
acceleration and HIC are biomechanical responses of the dummy. The
interpretation of head injury risk was con ducted using logistic probability
functions. Skull fracture and brain injury data for forehead impacts have been
used to calculate two-parameter functions for the probability of serious injury
during frontal impact [25]. The logist function data
for the Hybrid III 50th-percentile male was scaled using the procedure of Mertz
et al. [19] to the Hybrid III 5th-percentile female.
Unfortunately, no similar in jury risk functions are available for lateral
impact of the head.
Because of the lighter head and softer compliance of the female dummy neck,
the relative accelerations of the female dummy should be higher than those of
the male for comparable impact conditions. In addition, the tolerance to head
acceleration and HIC is also greater. This is consistent with the re cent
tolerance information from Mertz [18], which
indicates a tolerance of HIC = 1,000 for the midsize male and HIC = I,113 for
the small female dummy.
The logist relationship is a two-parameter sigmoidal function relating the
probability of injury p(x) = 1/[1 + exp(a - bx)] to a biomechanical response x,
where a and b are logist parameters. For a frontal forehead impact of the
50th-percentile Hybrid III male dummy, a = 2.048 and b = 0.0017 for HIC, and a =
2.859 and b = 0.0112 for peak resultant acceleration [32]. Scaling the risk parameters for the 5th-percentile
female gives a = 2.048 and b = 0.00153 for HIC, and a = 2.859 and b = 0.01006
for peak resultant acceleration.
RESULTS
Forehead impact responses
Figure 2 provides
the average peak acceleration and HIC based on the resultant acceleration for
the various baseball impacts on the forehead of the Hybrid III female dummy. Figure 3 shows the resultant head acceleration with the
official league hardball, reduced injury factor (RIF) baseballs, and softer
baseballs for the forehead impact. The softer base balls reduce the amplitude of
acceleration and in crease the impact duration from about 1 to > 2 ms. Figure 4 shows the resultant neck force for the same
tests.
The average peak head acceleration was 296 g and HIC was 434.5 for the
official hardball. The softer baseballs resulted in progressively greater
reductions in acceleration and HIC. The Worth RIF 10 had a 35.1% lower peak
acceleration at 192 g and a 46.0% lower HIC at 234.5, the RIF 5 had a 54.6%
lower peak acceleration at 134.5 g and a 69.5% lower HIC at 132.5, and the RIF I
had a 72.3% lower peak acceleration at 82.0 g and a 86.3% lower HIC at 59.5.
There were even lower values with the other soft baseballs. Thus, the RIF
baseballs resulted in a 35.1-72.3% reduction in head acceleration and a
42.0-86.3% reduction in HIC when compared to the official hardball. The lower
responses are statistically significant.
Fig. 2. Peak resultant head acceleration (dark bars) and head injury
criteria (HIC) (light bars) for forehead impacts by various baseballs. The data
on ProBall have been plotted twice to facilitate comparisons among the data
plots. RIF, reduced injury factor baseball.
Fig. 3. Resultant head acceleration (Acc) for an official hardball, three
reduced injury factor baseballs (RIF), and four softer baseballs for a forehead
impact.
Fig. 4. Resultant neck load for an official hardball, three reduced
injury factor baseballs (RIF), and four softer baseballs for a forehead impact.
Between the eyes impact responses
Figure 5
plots the average peak acceleration and HIC values for the various baseballs and
for frontal head impacts directed between the eyes. This impact results in
loading a portion of the nose and, therefore, involved a greater thickness of
the skin that covers the metal skull. The responses are generally lower than for
the forehead impacts.
Fig. 5. Peak resultant head acceleration (Res.Acc.,dark bars) and head
injury criteria (HIC) (light bars) for frontal imapcts between the eyes by
various baseballs. RIF, reduced injury factor baseball.
Right-side temple impact responses
Figure 6
shows the average peak acceleration and HIC values for the various baseballs and
for a right side head impact on the temple region. The acceleration and HIC
results are greater than observed in the frontal head impacts. This indicates a
greater risk of injury because of the lower tolerance to lateral head impact.
However, injury risk data are not available to evaluate the accelerations. Figures 7 and 8 provide examples of
the time histories of resultant head acceleration and neck load for various
baseballs.
Fig. 6. Peak resultant head acceleration (Res.Acc.,dark bars) and head
injury criteria (HIC) (light bars) for lateral impacts on the right temple by
various baseballs. RIF, reduced injury factor baseball.
Fig. 7. Resultant head acceleration (ACC) for an official, three reduced
injury factor (RIF), and four softer baseballs for a temple impact.
Fig. 8. Resultant neck load for an official, three reduced injury factor
(RIF), and four softer baseballs for a temple impact.
Injury risk assessment for head impacts
Figure
9 shows the probability of injury for the forehead impacts. It was developed
using the logist function and parameters, which convert biomechanical responses
(peak acceleration and HIC) to likelihood of injury. Based on HIC, the risk of
in jury varies from 11 to 20% for serious life threatening head injury of
Abbreviated Injury Scale (AIS) 3 + severity. The softer baseballs reduce
potential injuries. The RIF 10 represents a 22.0% reduction in risk, the RIF 5 a
32.0% reduction, and the RIF I a 38.0% lower risk in comparison to the official
league hardball. However, the softer balls had only a 4.4-7.6% incremental
difference in risk from the hardball.
Based on acceleration, the risk was 53.0% with the official league hardball.
The softer baseballs had lower risks of head injury: 46.6% lower with RIF 10,
65.7% lower with RIF 5, and 78.1% lower with RIF 1. This represents a 24.7-41.4%
lower increment in risk. The same injury probability calculations can be made
for the temple impacts. This is done for qualitative comparison only, because
the logist functions were developed for a frontal impact. On this basis, the
official hardball has a 51.6% risk, and the RIF I a 14.8% risk based on HIC.
These higher risks are consistent with the greater head accelerations and are in
part related to a thinner head skin covering the temple region.
Fig. 9. Risk of serious head injury calculated using logist functions,
which convert head injury criterion (light bars) (HIC) from biomechanical tests
to probabilities of injury by various baseball impacts on the forehead. RIF,
reduced injury factor baseball. The dark bars are head injury risks based on
resultant head acceleration.
DISCUSSION
The estimated risk of serious head injury using HIC for the forehead impacts
of the Hybrid III female dummy ranged from 20% for the official hard ball to 11%
for the softer baseballs. The RIF balls had an injury risk of 12.4-15.6% for a
forehead impact based on HIC and 11.6-28.3% based on head acceleration. The risk
of injury would be higher for the temple impacts. It ranged from 51.6% for HIC
and the hardball to about 11.8-28.7% for the softer balls, although this
projection is a conservation estimate based on frontal risk functions, since
there are a lack of injury data for lateral head impacts.
Based on the promotional literature from Heald [4], the risk of injury predicted from the NOCSAE
procedure is essentially 0 with RIF 1, 2% with RIF 5, and 10% with the RIF 10
baseball (Fig. 10); it is 80-99% with Major League and
popular Youth League baseballs. The Hybrid III indicates higher risks with the
softer baseballs and lower risks with a regular hardball. It is unclear at this
time which of the predictions is more representative of actual injury risks to
children, but the benefits in safety are probably fewer than previously
claimed.
The Hybrid III dummy has been widely accepted as the most advanced device for
impact testing. It was designed for direct head impacts on the fore head and has
been widely used for both hard and compliant head impacts. In addition, the
biofidelity of the Hybrid III neck structure, torso, and extremities should
provide a more realistic whole-body response. The whole body indirectly affects
head in jury assessment. Finally, most of the cadaver impact tests used to
determine injury risk functions were tests used to define the biofidelity of the
Hybrid III. While the biofidelity of the Hybrid III dummy has been established
for frontal impact on the forehead, more recent evaluations also demonstrate
biofidelity in head and neck responses for side impacts. The dummy is used for
injury assessment of pillar, rail, and windshield impacts in auto motive safety
tests and has been used in impacts of objects with sharp curvature and a wide
range in compliance.
The analysis of injury risk for the frontal forehead impacts of the Hybrid
III dummy has the greatest validity. The current probability functions relating
head acceleration and HIC to injury are limited to that type of loading
environment. However, we would have preferred to address the responses for the
temple impacts with side-impact head injury data and logist functions, since
lateral impacts are the greatest cause of injury for the unprotected head. An
analysis of lateral impact responses would show higher risks of injury, but the
influence of a thinner skin covering the skull and lack of realism in mimicking
compliance of the skull have an unknown influence on the tests. There are
probably two reasons for the higher incidence of human in jury for lateral
impacts: the individual is more likely not to see the ball coming from the side
and is less able to avoid the impact, and the skull is more compliant and less
protective of brain injury for lateral blows.
While the NOCSAE procedure may be an effective approach to evaluating head
protection in helmet impacts, there is minimal validation of the
head-neck-trolley model for direct head impacts. In contrast, the Hybrid 111
dummy has received inter national acceptance for its biofidelity of head and
neck impact. While this is true for a variety of impact conditions including
hard and compliant inter faces, there has been no validation of the dummy for
baseball impacts on the head. However, neither model has been validated for
impacts resulting in significant skull deformation in a lateral impact.
The WSU head impact procedure developed by Hodgson et al. [5,6] involves a dummy skull,
thin superficial skin, and a rigid neck structure. These parts are attached to a
trolley that responds indirectly to the head impact. This test method was
originally developed to assess the level of safety provided by helmet use in
baseball. Although it was later adopted in the NOCSAE helmet impact standard,
there may be important differences in simulating human responses for direct
baseball impacts. This may contribute to some of the differences in injury
predictions with the NOCSAE and the Hybrid III dummy.
The NOCSAE head-neck-trolley system has the neck rigidly connected to the
trolley and a skull in the head that deflects little during baseball impact.
Rather, the skin bottoms out with a hardball impact, the skull remains stiff,
and high accelerations occur. The effects of skin bottoming were seen in our
Hybrid III temple impacts with the official hardball. The high accelerations are
then related to a high probability of injury, which in turn are described as
related to a focal brain injury by skull deflection. Since the WSU and Hybrid
III heads don't deflect significantly during impact, the ensuing analysis of
data and injury risk interpretations are not tied to a biofidelity in dummy
response or mimicking of the underlying head injury mechanism. The rigidity of
the WSU neck may also involve an unrealistic mass from the trolley in the head
impact response in some long-duration situations, but the neck loads in the
Hybrid III female were typically of short duration and in the range of 70-350 lb
(300-1,500 N).
To our knowledge, this is the first study to report neck and contact loads
during baseball impact of the head. The neck loads closely followed the time
history of the head acceleration. Higher loads in the forehead impacts are
primarily related to the oblique orientation of the neck and fore-aft attachment
of the head to the neck. This is borne out by the higher z axis component of
neck load for the forehead impacts. During lateral impact, the primary component
of response was in the lateral (y axis) direction. Using the requirement for
equilibrium, the force of impact can be estimated by adding the inertial load
due to head acceleration to the measured neck load. This calculation results in
higher loads for the temple impacts.
Figure 11 shows the force of baseball impact for the
forehead and temple tests. The peak force ranged from about 2.5-18.0 kN
(550J.,000 lb). The major component of force is from head inertia as the neck
loads were only 6-8% of the total temple impact force. In contrast, neck loads
were 12-29% of the total force for the forehead impacts. The fraction of peak
load related to neck resistance in creased with the softer baseballs, which had
longer impact durations (Fig. 12). The RIF balls
provided a 35.6-72.2% reduction in peak impact force from that of the official
hardball. However, the official hardball resulted in such high-impact forces
that the biofidelity of the test dummy head and skin must be considered. The
forces are largely related to bottoming of the head-skin and interaction with
the hard metal skull. In a comparable situation with a human impact, deflection
of the skull and potential fracture would be expected to limit the maximum load.
Both the Hybrid dummies and the NOCSAE procedure would need further
qualification as realistic tools under extreme impact loading. Nevertheless, the
high loads and head acceleration are related to a greater risk of head injury as
indicated in the assessment of risk.
The high-speed videotape (Fig. 13) showed a larger
contact area with the softer baseballs than with the official hardball. A larger
contact area produces a proportionately lower pressure of ball con tact. The
lower force and pressure with the softer baseballs significantly reduce the risk
of skull fracture. While this has an obvious benefit, there is a need to
consider a secondary injury mechanism related to brain deformation from closed
head acceleration. The evaluation of neural injury risk would require the
measurement of head rotational acceleration in addition to the translational
measurement made in this study. Gennarelli et al. [3]
have demonstrated the importance of head rotational acceleration in the
production of neural injury in lateral impacts. In addition, the hardball might
cause more glancing blows of the head than the softer baseballs, which have a
greater compliance, develop a larger area of contact, and may increase
interactions with the head.
Qualitative observations of the high-speed video tapes provide an estimate of
the rebound velocity for the lateral impacts at about a quarter of the impact
velocity. The softer baseballs showed a lower rebound speed, but the high-speed
videotape at 500 frames/s provided only general documentation of the tests. More
exact information on rebound speed would have to be obtained using 16-mm
high-speed cine and a faster frame rate than possible with the high-speed
videotape. This information would help determine the energy transfer to the head
by base ball impacts.
While the NOCSAE standard developed by Hodgson at WSU has been in effect
since 1978 and undergone seven revisions through 1991, the head neck-trolley
test procedure has been fundamental to the evaluation of helmet safety. With the
development of softer baseballs, which are advertised as improving safety for
children playing baseball there is a need for an objective test to quantify the
safety provided by the newer equipment. While the use of the NOCSAE standard was
approved for evaluating the safety of baseballs in 1989 and may provide an
initial estimate of safety effectiveness, a validated procedure is required to
measure the underlying risk of direct impact by baseballs, soft balls, hockey
pucks, and other blunt objects used in sport. The 5th-percentile Hybrid III
female dummy may provide that device and procedure. Only further testing will
determine the appropriate certification of safety equipment, particularly as the
procedure needs validation for closed head injury of the brain and skull
deflection for temple impacts by a baseball.
Fig. 10. Probability of injury with various baseballs from promotional
material distributed by the Worth Company (1991) and from the Hybrid III female
tests in this study. The risks for temple impacts were calculated using the
forehead impact logist function data and are plotted here as a conservative
estimate of the risk. RIF, reduced injury factor baseball.
Fig. 11. Peak impact force by various baseballs for forehead (light bars)
and temple (dark bars) loading at 60 mi/h. RIF, reduced injury factor baseball.
Fig. 12. Head impact force for an official hardball, three reduced injury
factor baseballs (RIF), and four softer baseballs for a forehead impact (A) and
a temple impact (B).
Fig. 13. Images from the high-speed videotape of various baseball impacts
on the temple of the Hybrid III female dummy.
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This article was published as:
"Analysis and Comparison of Head Impacts
Using Baseballs of Various Hardness and a Hybrid III Dummy"
Clinical
Journal of Sports Medicine
Vol. 3, 1993, pp. 217-28.
Viano DC,
McCleary JD, Andnejak DV, Janda DH
It is possible to order a copy of this article.
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