Finite Element Modeling of an Innovative Headband1127992/FULLTEXT01.pdf · Finite Element Modeling...

DEGREE PROJECT IN MEDICAL ENGINEERING, SECOND CYCLE, 30 CREDITS

STOCKHOLM, SWEDEN 2017

Finite Element Modeling of an Innovative Headband

A study about the headband design and its safety properties

GIULIANA MOZZI

KTH ROYAL INSTITUTE OF TECHNOLOGY

SCHOOL OF TECHNOLOGY AND HEALTH

Finite Element Modeling of an Innovative Headband

A study about the headband design and its safety properties

GIULIANA MOZZI

Finita Element Modellering av ett innovativt huvudband

En studie om huvudbandets design och dess skyddsegenskaper

GIULIANA MOZZI

Master of Science - Thesis in Medical Engineering, 2017 Advanced level (second cycle), 30 credits

Supervisors: Svein Kleiven and Hans von Holst Reviewer: Svein Kleiven Examiner: Mats Nilsson

KTH Royal Institute of Technology

School of Technology and Health (STH), KTH SE -141 86 Flemingsberg, Sweden

http://www.kth.se/sth

3

Abstract

One of the most recent challenges faced worldwide is the so-called “ageing population”

phenomenon. Besides the disabilities and illnesses that the elderly is prone to, traumatic

brain injuries represent a common and potentially fatal factor that greatly affects this

population. Other subjects are however also likely to experience traumatic brain injuries i.e.

people affected by some kinds of diseases and disorders such as epilepsy and dementia that

are known to have a high risk of falling. Furthermore, sports accidents can also lead to

traumatic brain injuries, for example cycling. Current solutions and devices for head injury

prevention are limited and from a design point of view there is space for improvements. The

focus of this master thesis is to model an innovative headband meant to be worn around the

scalp during various daily life activities and in multiple situations. Its function is to prevent

the wearer from traumatic brain injuries. The headband model is a finite element model,

created in LS-PrePost. Subsequently, simulations are performed in LS-Dyna to replicate real

life head impact scenarios. Different parameters and features of both the impacting

condition and the headband are tested for creating the optimal headband structure for head

injury prevention.

5

Acknowledgements

I would first like to thank my supervisors Dr. Hans von Holst and Dr. Svein Kleiven for giving

me the opportunity of being part of this very interesting project and for their guidance during

this semester. Their professional suggestions and ideas were always extremely helpful and

inspiring.

I would also like to thank Xiaogai Li, who provided expertise that greatly assisted my thesis

report, and Hildi, Isotta and Silvia for their precious feedback and daily support.

Finally, words of sincere and deep gratitude go to my family, that has always supported me

no matter what and has made this two-year journey possible.

7

Table of Contents

1 Introduction ...................................................................................... 1

2 Aim .................................................................................................... 3

3 Methods............................................................................................. 5

3.1 Headband structure ....................................................................................................... 5

3.2 Pin configurations ......................................................................................................... 6

3.3 Headband configurations .............................................................................................. 7

3.4 Impact scenarios ........................................................................................................... 8

3.4.1 Falls from a standing position ................................................................................ 8

3.4.2 Fall from a bicycle ................................................................................................... 9

4 Results ............................................................................................. 11

4.1 Falls from a standing position ...................................................................................... 11

4.1.1 Influence of rubber material properties ................................................................. 11

4.1.2 Influence of connecting elastic beams material properties .................................. 13

4.1.3 Influence of impact locations ................................................................................ 13

4.2 Falls from a bicycle ...................................................................................................... 17

5 Discussion ....................................................................................... 21

6 Conclusion ...................................................................................... 23

7 References ....................................................................................... 25

Appendix A ......................................................................................... 27

Appendix B ......................................................................................... 55

Appendix C ......................................................................................... 57

9

List of Abbreviations

DAI Diffuse Axonal Injuries

FE Finite Element

HB Headband

KTH Kungliga Tekniska Högskolan

LSTC Livermore Software Technology Corporation

SFC Skull Fracture Correlate

TBI Traumatic Brain Injuries

WHO World Health Organization

1

1 Introduction

At the present, one of the most relevant issues that humanity faces worldwide is the

increasing age of the population. This phenomenon has arisen from great achievements in

terms of technology, socioeconomic development and public health. However, this also

introduces a challenge to the entire society in terms of guaranteeing the security and well-

being of the elderly. The World Health Organization (WHO) has with statistical reports

estimated an increase in the number of people aged 60 years and older, from 688 million in

2006 to about two billions by 2050 [1].

Falls are known to be the leading cause of injuries resulting in severe influence on the

physical and psychological well-being of the elderly. According to WHO, falls are the leading

cause of traumatic brain injuries (TBI), with 32-42% of the population aged 70 years and

older subjected to a fall every year [1]. Between 2002 and 2006, 595 095 cases of TBI related

to falls were registered in the USA [3]. This highlights the alarming figure of the drastic

increment in the amount of people affected by injurious falls, as well as their consequences,

especially among the elderly. One of the consequences is increased costs, not only for the

affected individuals, but also for the entire society.

Other groups of the population can also be pointed out to have a high risk of injurious falls

i.e. people suffering from diseases or disorders, or people engaged in sports or activities such

as cycling. Dementia is a typical example of a disease that accounts for a higher rate in falls

among the elderly [4]. The reasons primarily lie in the chronic deterioration of the cognitive

and motor functions [5], with consequent issues in ambulation and limited range of motion.

These factors greatly influence the capability of having control over the body to withstand a

fall. By 2030 and 2050, the population affected by dementia is estimated to reach 75.6

million and a tripled increase, respectively [6]. With regards to cycling, it has been associated

to 85 389 cases of TBI treated at the emergency room in 2006, only in the USA [7]. Some of

the most common head injuries are contusions, hematomas, concussions, diffuse brain

injuries (DAI) and skull fractures. Depending on the severity of the lesion, the outcomes can

vary from mild conditions to fatal events.

Prevention plays a key role in limiting not only the incidence and severity of head injuries

but also the costs associated with injurious falls. Some of the typical approaches that are

present in today prevention strategies consist in the adoption of walking aids, performance

of strength and balance trainings, introduction of vitamin D and calcium to the diet etc. [8].

However, remarkable results have not been achieved by these methods and importantly,

they do not provide any protection at the moment of an impact. Researchers have therefore

increased their focus on this aspect with some new emerging solutions on the market. For

instance, highly engineered materials and devices have been designed to absorb most of the

energy during impacts so that brain and skull get subjected to milder lesions. Bike helmets

are examples of outstanding solutions, known to greatly reduce the risk of TBI in bike

accidents [9]. Functionality, wearability and comfort of those on the market are however not

optimal. Hövding [10] is an example of an inflatable airbag-like helmet that provides an

innovative solution. Its limitations concern high costs, lack in the accuracy of the activation

mechanism for inflation and the impossibility to reuse the product after it being expanded.

Protective headbands, designed for different settings such as in the car or during a soccer

match, can also be found on the market [11]. But the range of available products seems rather

2

limited. Furthermore, no testing data was found in the literature on the effectiveness of these

headbands. This suggests that there should be space for new designs on the market. Finally,

energy absorbing materials incorporated in floor systems represent the edge in technology

related to fall prevention. More specifically, the flooring system created at Kungliga Tekniska

Högskolan (KTH) by Prof. Svein Kleiven and Prof. Hans von Holst allows a reduction of 60-

75% and 39-66% for respectively impact forces and accelerations [12].

Combining the effectiveness of the highly protective properties exhibited by these flooring

systems, and normal headband designs, in terms of shape, wearability and comfort, could

give rise to a promising product for protection. The proposed headband could not only come

as an alternative to the traditional bike helmet, but also it could serve as a product with

potential medical applications. It could be worn by people who are likely to fall and hit their

head i.e. the elderly or patients who are affected by seizures. The multiple functions and

situations in which this product could be worn make it relevant for safety issues related to

the healthcare sphere, as well as to a much larger scale that comprehends everyday life

scenarios. In order to study the safety properties of such a headband, a numerical model can

be created and different impact simulations can be performed, investigating its efficacy

during injurious fall situations. This master thesis aims at performing these tasks to provide

consistent evidence of the effectiveness of this innovative headband, which could potentially

fulfill the modern market demand.

3

2 Aim

This master thesis aims at investigating the protective properties of an innovative headband

that finds its application in prevention and attenuation of head injuries. In order to

accomplish this goal, a virtual model of the headband was created using a pre- and post-

processor software and then impact simulations were performed by creating different

impact scenarios between the ground and the KTH head model while wearing and not

wearing the headband. The different configurations of the structure and material properties

of the headband were explored to achieve the optimal outcomes in terms of safety and design

requirements. In particular, the effectiveness of the product was investigated by simulating

falls from a standing position and from a bicycle. In the latter case, the parameters of the

impact replicate the same conditions simulated in a previous finite element (FE) study [13]

in which the efficacy of a bike helmet was analyzed. Finally, the results obtained from the

simulations were evaluated against thresholds from the literature to predict the occurrence

and severity of brain and skull injuries. The present study is particularly useful to

demonstrate not only the possibility of using this new product as a potential alternative to

modern bike helmets, but it also provides evidence of a cutting-edge device that can bring

significant benefits to the entire society.

5

3 Methods

The innovative headband model and the analysis of the results was performed using LS-

PrePost® V4.3 [14], an advanced pre- and post-processor by Livermore Software

Technology Corporation (LSTC). Finally, the impact simulations were performed using the

software LS-DYNA (version R8.0.0, revision 95309) [15], by LSTC. The following

paragraphs describe the different steps and decisions that were made to create the virtual

model of the innovative headband (HB), from its core structure and material properties

modeling to the testing of its efficacy under impact simulations that replicate real-life

scenarios of typical injurious falls.

3.1 Headband structure

The material used to create the core part of the headband is the same that is in use in the

KTH shock absorbing floor and that has been modeled in a previous Master thesis [16]. One

pin was isolated from the floor (Fig. 1) and its structural features were modified in order to

obtain four different configurations as described in more detail in the “Pin configurations”

section. The modified stud was subsequently reflected several times to create a string

composed by several lines of pins for a total headband width of 4 cm as shown in Fig. 2.

Fig. 1 - KTH energy absorbing system and extracted pin.

Fig. 2 - Side and top view of the final components of the headband.

Bottom layer

Overlying layer

Pins

layer

Front Rear

TOP VIEW

Inner

layer

SIDE VIEW

4 cm

Pin layer

6

The string is composed of solid elements made of styrene butyrenic rubber and the material

used to model this component is *181-SIMPLIFIED_RUBBER/FOAM. The string was

manually positioned around the head and additional manual adjustments such as rotations

and translations of the pin nodes were required to prevent computational errors due to

negative volumes and initial penetrations as depicted in Fig. 3. Moreover, an inner rubber

solid layer was introduced in the interface head-pins to prevent

the pins from buckling. Furthermore, this layer allowed to avoid

computational instabilities and its thickness was varied from 1

mm to 3 mm to study which structure was more effective from a

protective point of view. Pins and solid layer are made of the same

rubber material that can be found in the floor but its original

stiffness was modified due to large deformations in the HB during

the impact simulations. Moreover, a third rubber layer was

originally introduced at the bottom of the pins and subsequently

removed since it did not bring any relevant benefit to the

previously mentioned structure. Finally, to create an optimal

fitting to avoid gaps between the head and the HB, a set of

concentric forces was applied around the HB and its final

structure is depicted in Fig 2.

Lastly, eight different configurations were created depending on the pins and HB features.

More details about these aspects are provided in the following paragraphs. In order to

evaluate the best configuration of the HB in terms of safety properties, the analysis of the

results was focused on the evaluation of the damages occurred in the brain and in the skull

bone. In particular, the peak value of the first principal Green-Lagrange strain has been

found to be well correlated with brain injuries [17] whereas HIC15 and peak resultant

acceleration of the center of gravity of the head can predict the risk of skull fractures

according to the curves developed by Mertz et al. [11] and Chan et al. [18]. Accelerations and

velocities were obtained using SAE 300 filter.

3.2 Pin configurations

The original structure of the KTH energy absorbing floor was modified and thus four

different configurations of the HB pins were obtained. Fig. 4 addresses the parameters that

were altered where D stands for the distance between the studs, L their length, and Φ their

diameter. The values chosen for the mentioned parameters are presented in Tab. 1. The layer

that connects the pins was modified by changing its thickness from 3 mm to 5 mm to study

the optimal structure.

Fig. 3 - Initial penetrations causing computational instabilities.

7

3.3 Headband configurations

Starting from an intact structure, the HB layers were divided into four pieces as illustrated

in Fig. 5. The connective parts are made of elastic rubber beams

(*MAT_LINEAR_ELASTIC_DISCRETE_BEAM) which were studied in terms of stiffness

and presence/position by means of radial lateral impacts at vi = 5.9 m/s. Consequently, for

every intact configuration the 4-pieces counterpart was modeled. For sake of clarity, the

several HB structures are addressed by a capital letter for intact configurations (A, B, C, or

D) and by a lowercase letter otherwise (Tab. 2). In particular, for type a, the stiffness was

augmented by one order from the baseline values whereas the position and their presence

was varied as illustrated in Fig. 6. The total number of elements that form the different final

HB structures (pins layer + solid rubber layer / + connecting beams) are respectively 28 000,

27 185, 36 209, 34 603, 44 388, 42 400, 112 801, and 110 101 for A, a, B, b, C, c, D, and d.

Approximately, the weight of the HB varies in the range of 176-272 g.

TYPE L [mm] D [mm] Φ [mm]

A + ++ ++

B ++ ++ ++

C +++ ++ ++

D + + +

TYPE

Intact 4-pieces

A a

B b

C c

D d

SIDE VIEW TOP VIEW

Fig. 4 - Side and top view of the pins showing the features that have been modified for the different configurations.

Tab. 1 - Values related to the different configurations of the studs. For confidentiality reasons, the values have been masked.

Front Rear

Tab. 2 - Types of HB structures.

Fig. 5 - Intact and 4-pieces HB configurations.

8

Fig. 6 – (A) Beams present only between the pin pieces; (B) Additional beams connecting the bottom of the solid layer parts; (C) Additional beams connecting the top of the solid layer parts.

3.4 Impact scenarios

Simulations representing fall impacts under two main scenarios were performed – falls from

a standing position and from a bicycle.

3.4.1 Falls from a standing position

The head model used was developed at KTH and more details about it are provided in the

Background chapter. The locations of impact were the side and the rear of the head, which

were found to be the most common areas for this kind of fall [19]. In particular, two cases

were taken from a study conducted at the NHMRC Road Accident Research Unit at the

University of Adelaide in which brain injury patterns related to fatal falls from a standing

position were collected from the victims [19]. For both cases, the initial velocity was

estimated based on the assumption that the victim had fallen from his/her own height,

reaching an initial impact velocity of 5.9 m/s (Case I, Fig. 7). The same impacts were also

tested at a speed of 3.5 m/s (Case II, Fig. 7). Furthermore, the material properties of the

rubber components were investigated by varying the stiffness (Curve 1, see Appendix B) and

the thickness of the HB layers while keeping the same structure (type A). In particular, the

original Curve 1 was changed by increasing and decreasing the stiffness scale factor by one

order. In the model, the colliding parts are interacting by contact type

*AUTOMATIC_SURFACE_TO_SURFACE and the coefficient of friction used between

head-HB and HB-ground equals 0.45. Moreover, the ground was modeled like the rigid

asphalt realized by Fahlstedt et al. [13].

A B C

9

Fig. 7 - Lateral and occipital radial impacts against the asphalt while wearing the HB at vi=5.9 m/s and vi=3.5 m/s.

3.4.2 Fall from a bicycle

The case analyzed in this section replicated the same real impact scenario taken into account

within the study performed by Fahlstedt et al. [13] and in which the effectiveness of wearing

a bike helmet was tested. Same velocities, accelerations, impact angles and impacting rigid

asphalt of the mentioned study were tested with the additional presence of the HB (Fig. 8).

In particular, the impact is characterized by a resultant velocity of 5.3 m/s and a resultant

acceleration that equals 4.7 rad/s2. More details about this case can be found in previous

studies [18, 19].

Fig. 8 - Lateral oblique impact while wearing the HB in bicycle accidents.

Case I: vi = 5.9 m/s

Case II: vi = 3.5 m/s

X

Z

Y

Z

Z

Y

Z

X

vres = 5.3 m/s

wres = 4.7 rad/s2

Case I: vi = 5.9 m/s

Case II: vi = 3.5 m/s

11

4 Results

4.1 Falls from a standing position

4.1.1 Influence of rubber material properties

In Tab. 3, the values show different outcomes obtained from a radial impact at vi = 5.9 m/s.

Base represents the connective layer at the bottom of the pins, Solid refers to the top solid

rubber layer, and the number next to these headings stands for the thickness of the layer in

mm. The most relevant results are obtained when adopting Curve 1 (see Appendix B), a base

of 5 mm and a solid 2 mm: brain strain and resultant linear acceleration are reduced by

respectively 13% and 45% when compared to the bare head case.

Tab. 3 - Outcomes for different cases, HB structure and rubber material stiffness (Curve 1). The numbers that follow the Base and Solid captions represent the thickness of the solid and base layers in mm.

*Averaged acceleration over HIC15.

CURVE 1 N0 HB Base 3

Solid 1

Base 3

Solid 2

Base 5

Solid 1

Base 5

Solid 2

Base 5

Solid 3

Peak 1st principal strain in the brain

Baseline

Stiffer

Softer

0.62

0.62

0.62

0.55

-

-

0.55

0.64

0.61

0.54

0.63

0.61

0.54

0.62

0.63

0.57

-

-

Risk of concussion based on risk curve from Kleiven [17]

Baseline

Stiffer

Softer

1

1

1

0.97

-

-

0.97

1

0.99

0.97

1

0.99

0.97

1

1

0.99

-

-

Peak von Mises stress in the skull [MPa]

Baseline

Stiffer

Softer

53.44

53.44

53.44

42.65

-

-

42.35

51.63

45.16

42.13

52.15

44.97

36.55

44.85

43.80

35.62

-

-

Skull Fracture Correlate (SFC)* [g]

Baseline

Stiffer

Softer

342.57

342.57

342.57

273.44

-

-

266.14

318.26

319.76

256.70

310.55

331.10

204.24

286.04

288.63

198.75

-

-

Risk of skull fracture based on risk curve from Chan et al. [18]

Baseline

Stiffer

Softer

1

1

1

0.94

-

-

0.94

1

1

0.91

1

1

0.75

0.96

0.96

0.71

-

-

12

A stiffer curve (Curve 2, see Appendix B) was considered and the same cases from Curve 1

were replicated. The HB exhibits less deformable behavior when stiffened (Fig. 9) and higher

values as concerns the 1st principal strain in the brain, von Mises stress in the skull and

resultant linear acceleration as shown in Tab. 4. In particular, stiffening the material curves

by one order results in strains higher than when not wearing the HB.

Fig. 9 – Visual stiffness testing results from case Curve 1- Baseline; Base 5, Solid 2 and Curve 2 – Baseline; Base 5, Solid 2 at the max. deformation of the HB during the impact.

Tab. 4 - Outcomes for different cases, HB structure and rubber material stiffness (Curve 2).

*Averaged acceleration over HIC15.

** Based on the SFC.

CURVE 2 No HB Base 3

Solid 1

Base 3

Solid 2

Base 5

Solid 1

Base 5

Solid 2

Base 5

Solid 3


Baseline

Stiffer

Softer

0.62

0.62

0.62

0.56

-

-

0.56

0.65

0.57

0.54

0.65

0.65

0.59

0.63

0.60

0.59

-

-


Baseline

Stiffer

Softer

1

1

1

0.98

-

-

0.98

1

0.99

0.97

1

1

0.99

1

0.99

0.99

-

-


Baseline

Stiffer

Softer

53.44

53.44

53.44

38.35

-

-

38.38

53.08

44.06

39.49

52.54

52.64

42.88

49.86

43.92

38.75

-

-

Skull Fracture Correlate (SFC)* [g]

Baseline

Stiffer

Softer

342.57

342.57

342.57

277.63

-

-

256.78

341.37

313.10

271.95

331.52

317.50

205.12

336.80

262.14

197.26

-

-

Risk of skull fracture based on risk curve from Chan et al.** [18]

Baseline

Stiffer

Softer

1

1

1

0.95

-

-

0.91

1

1

0.94

1

0.99

0.74

1

0.94

0.70

-

-

13

4.1.2 Influence of connecting elastic beams material properties

For this study, ‘Curve 1 – Baseline’ and ‘Base 5, Solid 2’ structure were adopted. The presence

of the elastic beams (Case B, B stiffer and C) affects the outcome values lowering the strain

but increasing the resultant linear acceleration and peak von Mises stress in the skull (Fig.

10). Moreover, the position of the beams connecting the solid layer parts influences the

results. The lowest strain is reached when connecting the stiffer beams to the bottom sides

of the rubber solid layer (Case B stiffer). When compared to the bare head impact, this

configuration allows to reduce the strain by almost 61%.

Fig. 10 - Peak 1st principal strain (left) in the brain and peak von Mises stress in the skull (right) obtained using Curve 1 varying the stiffness, presence and position of the connecting elastic beams.

4.1.3 Influence of impact locations

The several configurations listed and described in the Methods section were tested on radial

impacts on the side and on the rear of the head at vi = 5.9 m/s and vi = 3.5 m/s.

For lateral crashes at vi=5.9 m/s, the 4-pieces configuration offers overall better protection

from brain strains than the corresponding intact structure: a reduces deformation by 44%

and 52% when compared respectively to A and no HB case (see Tab. 5). In general, 4-pieces

structures imply slightly higher stresses in the skull than in the intact shape but lower than

the bare head impact. In terms of strain in the brain, using the HB brings benefits in most

of the lateral impacts at high speed but the same conclusion cannot be drawn from the

occipital accidents results. However, remarkable lower values related to the von Mises stress

in the skull and translational accelerations can be observed in both lateral and occipital

impacts at high and low velocities. In order to protect the subject from concussion at high

speed falls, wearing a when a lateral impact occurs can reduce the risk of brain damages

from 100% to 64%. On the contrary, to protect the brain during occipital impacts, the present

configurations do not provide benefits from this point of view. However, the severity of

injuries related to linear accelerations and occipital falls can be minimized by wearing any

of the HB configurations. The rotational velocities and accelerations that characterize the

different impact scenarios are presented respectively in Tab. 6 and Tab. 7. Finally, at vi=3.5

m/s the HB exhibits higher protective properties for brain injury in lateral impacts with a

reduction of 47% regarding the strain. For occipital impacts, instead, benefits arise in the

lower stress values associated to skull fractures.

0,55

0,30,24

0,4

0

0,1

0,2

0,3

0,4

0,5

0,6

Peak 1st principal strain

Case A Case B Case B stiffer Case C

36,02

38,8738,22

38,89

34

35

36

37

38

39

40

Peak von Mises stress [MPa]

Case A Case B Case B stiffer Case C

14

Standing No HB A a B b C c D d


Lateral I

Lateral II

Occipital I

Occipital II

0.62

0.47

0.19

0.11

0.54

0.39

0.35

0.21

0.30

0.36

0.32

0.21

0.65

0.39

0.37

0.23

0.54

0.31

0.29

0.21

0.62

0.35

0.36

0.23

0.54

0.25

0.26

0.18

0.60

0.39

0.32

0.15

0.50*

0.30

0.44

0.21


Lateral I

Lateral II

Occipital I

Occipital II

1

0.95

0.27

0.12

0.97

0.86

0.80

0.32

0.64

0.81

0.70

0.32

1

0.86

0.82

0.40

0.97

0.69

0.65

0.32

1

0.80

0.81

0.40

0.97

046

0.50

0.25

0.99

0.86

0.70

0.19

0.96*

0.64

0.95

0.32


Lateral I

Lateral II

Occipital I

Occipital II

53.44

46.23

51.75

43.90

36.55

24.19

37.64

19.28

38.22

27.68

38.87

20.15

35.24

12.94

40.38

10.90

36.61

14.59

41.01

14.85

26.87

17.79

40.99

10.12

35.36

13.59

41.70

14.91

37.96

27.91

38.79

19.82

53.43*

33.05

40.65

20.41

Peak resultant linear acceleration (1) and (2) Peak HIC15 [g]

Lateral I(1)

Lateral II(1)

Occipital I (2)

Occipital II(2)

625.37

397.83

662.85

438.22

249.11

204.15

331.22

141.37

279.32

209.32

319.64

145.23

256.54

143.04

340.48

116.03

232.85

114.28

363.56

102.60

210.11

117.36

380.96

82.02

159.45

111.43

425.02

99.86

270.28

205.01

323.37

138.66

448.89*

236.71

340.84

139.17

Tab. 5 – Lateral and rear impacts from a standing position at different speeds (Case I: 5.9 m/s; Case II: 3.5 m/s) while wearing and not wearing the different HBs.

14

15

*Stiffness augmented by one order due to large deformations in the rubber layers.

** Based on the strain.

Risk of skull fracture based on risk curve from (3) Chan et al.** [18] and (4) Mertz et al. [21]

Lateral I(3)

Lateral II(3)

Occipital I(4)

Occipital II(4)

0.99

0.92

0.02-0.05

0.01-0.02

0.95

0.83

0.005-0.01

0.001-0.005

0.54

0.78

0.005-0.01

0.001-0.005

0.99

0.83

0.005-0.01

0.001-0.005

0.95

0.58

0.005-0.01

0.001-0.005

0.99

0.76

0.005-0.01

0.001-0.005

0.95

0.27

0.005-0.01

0.001-0.005

0.98

0.83

0.005-0.01

0.001-0.005

0.95*

0.54

0.005-0.01

0.001-0.005

15

16

No HB A a B b C c D d

Lateral I 49.14 44.60 20.00 61.08 51.88 56.74 51.11 53.82 40.01*

Lateral II 36.22 33.49 29.12 35.47 31.21 32.44 26.07 33.19 29.16

Occipital I 5.71 28.00 23.47 29.44 26.90 27.83 23.85 27.07 24.32

Occipital II 3.77 16.60 13.91 19.17 16.81 19.90 16.41 12.55 15.31


No HB A a B b C c D d

Lateral I 43.01 32.19 23.09 25.50 27.52 23.57 21.33 29.73 58.28*

Lateral II 28.08 12.17 13.35 13.18 10.73 9.14 9.69 13.95 15.64

Occipital I 24.25 21.12 24.62 19.76 15.73 13.93 19.19 24.83 24.74

Occipital II 14.35 8.28 11.28 7.56 18.96 7.16 10.43 6.40 9.51


Tab. 6 – Rotational velocity in rad/s for lateral and rear impacts at 5.6 m/s (Case I) and 3.5 m/s (Case II).

Tab. 7 – Rotational acceleration in krad/s2 for lateral and rear impacts at 5.6 m/s (Case I) and 3.5 m/s (Case II).

17

4.2 Falls from a bicycle

In bicycle accidents, the different structures of the HB provide a remarkable protection against

the risk of skull fractures. As highlighted in Tab. 8, the peak resultant linear acceleration drops

from 418.83 g to 130.84 g when wearing b. In all the configurations, the stresses applied to the

skull are reduced by more than half of the reference case value (No HB). In particular, a pattern

has emerged for lateral impacts: longer pins imply lower values in the von Mises stresses that

occur in the skull. In fact, in terms of skull stresses, C shows the lowest value as highlighted in

Fig. 11. This value reaches 16.72 MPa, which can be closely related to the outcome obtained when

wearing the helmet, 16 MPa. On the contrary, shorter pins and intact configurations are more

effective in reducing the strain in the brain for high speed cases. When wearing B, the lowest

value is obtained and it equals 0.30 which stands between the values related to no HB case

(0.43) and helmeted head case (0.24).

HB - configuration C No HB

0 MPa 80 MPa

Fig. 11 - Von Mises stress in the skull for bare head and head wearing C.

18

18

19

Bicycle No HB Helmet A a B b C c D d


Lateral 0.43 0.24 0.32 0.34 0.30 0.35 0.38 0.42 0.32 0.37


Lateral 0.92 0.41 0.73 0.75 0.66 0.76 0.85 0.92 0.73 0.82


Lateral 54.88 16.00 25.00 27.93 22.35 26.93 16.72 22.99 22.05 23.96

Peak resultant linear acceleration [g]

Lateral 418.83 149.00 182.24 176.14 141.11 130.84 132.48 144.64 168.92 160.83

Risk of skull fracture based on risk curve from Chan et al. [18]

Lateral 0.88 0.27 0.62 0.69 0.54 0.72 0.82 0.87 0.62 0.78

Tab. 8 – Lateral oblique impacts with bare head, while wearing a bike helmet and different HBs.

19

21

5 Discussion

The modeling of the HB represents the fundamental aspect that shapes this master thesis.

Every step and simulation was performed in order not only to create the model itself but also

to study its characteristics and features to obtain a virtual model capable of replicating real

life impact scenarios to predict head injuries and countermeasures to prevent them. Shaping

the HB around the head in a semiautomatic way by using initial velocities applied to the pins

string so that it would wrap around the scalp was a challenge that implied long simulations

so that the manual technique described in the Methods section was preferred. Additionally,

manual adjustments allowed having more control over the shaping process by constantly

checking penetrations among the several parts and overcoming the issue of negative volume

errors occurring in the initial simulations. Regarding the different types of pins, the first

three structures allow to compare one to each other to evaluate the influence of the pins

length. The fourth configuration was suggested from previous studies about the KTH floor

system since it showed good results in hip impacts. Finally, the intact HB configuration was

cut into 4-pieces and positioned around the head model so that the rubber part would be

located in the impact area. In this way, the connecting elastic beams were not directly

interfering between the head and the ground. Additionally, the HB was divided into 4-pieces

to satisfy manufacturing and design requirements. Producing smaller parts and connecting

them through elastic beams allows creating molds in an easy way and having a better fit over

the head since the structure can adapt to different sizes and stay tight around the scalp. To

evaluate the protection of different configurations of the HB, a few representative cases were

selected.

The choice of adopting an initial impact velocity that equals 5.9 m/s for the material study

relies in two main reasons: firstly, the targets of this device are characterized by two major

impact scenarios such as falls from a standing position and from a bicycle. In the latter case,

higher velocities and deformations occur so it seemed appropriate to focus on creating a

material capable of being sound and effective in such impacts. Secondly, the mentioned

velocity provides the starting point to analyze the worst-case scenario in falls from a standing

position in which a free fall from the person’s own height is taken into account. For instance,

the velocity of 5.9 m/s characterizes a free fall of the head from 1.80 m. Although free falls

do not occur often and the head does not represent the first impacting part of the body,

studies have proven that for instance the use of the hands has no protective efficacy among

the elderly due to the fact that motor reflexes and muscles are delayed and weakened [22].

From the results obtained in the initial material properties study, it can be stated that Curve

1 exhibits overall slightly lower values than Curve 2 as concerns strain in the brain and SFC.

Additionally, from a visual point of view, the first curve appears more deformable, thus

suggesting a better ability to withstand deformations and absorb shock energy. However,

both material curves show good properties of elastic shape recovering after the impact. A

further observation regards the different levels of stiffness of the curves. A softer behavior

shows the incapability of withstanding the impact since the pins deform greatly without

providing enough strength. On the contrary, stiffer materials do not provide the “cushion

effect” that can be observed in the previously mentioned cases and thus the head gives

22

evidence of more severe injuries. For the above-mentioned reasons, the material that was

chosen to start modeling the different HB configurations is Curve 1 – Baseline.

From the obtained results, the 4-pieces structures reduce strain values in the brain due to

the fact that the energy is transferred to the HB and used to deform the rubber rather than

the brain and skull. The efficacy of the HB reflects the highly energy absorptive properties

shown by the KTH floor system. The material used in the floor shows a reduction of 77% in

the translational acceleration (see Appendix C), whereas the modified material used within

the HB reaches a corresponding max. linear acceleration decrement of 74% (Case Lateral I

when wearing c - 4-pieces, L=iii, D=dd, Φ=ff). For lateral impacts, longer pins decrease the

von Mises stresses in the skull and, when applied in occipital accidents, they show lower

peak HIC15 and lower risk of skull fractures. The bending behavior of the pins allows not only

having smoother interactions between the colliding parts but it also provides an effective

way to dissipate the shock energy by transferring the deformations from the head to the HB.

Although the use of the HB provides overall protection to the brain and the skull, the rear

impacts show an increment in the brain strain when wearing the HB due to the rotational

component related to the velocity (see Tab. 6). Thus, it can be suggested that different impact

locations require different HB characteristics to achieve optimal results but further studies

must be performed.

The present master thesis provides a starting point to study the optimal characteristics that

the proposed innovative HB should integrate. From the material to the structure, every

component has a fundamental role in the obtained outcomes. The simulations that were

performed allow to make a reasonable comparison among the different configurations that

were taken into account. However, combining different features and changing values and

parameters is a very laborious work that cannot aim at investigating all the possible

scenarios due to the intrinsic variability of fall cases. This aspect represents one of the limits

of this thesis which should be taken as an initial step to perform more targeted simulations

that could gather the best results achieved so far and investigate different parameters to

obtain optimal results. Additionally, the use of the single head model instead of the total

human body model could influence the outcomes although from the literature it emerged

that it is common practice to perform and use just the head model to simulate and analyze

head impacts. Furthermore, as already mentioned, more in-depth studies could focus on a

lower number of HB configurations and vary the parameters of interest (i.e. friction

coefficients, material properties etc.) to achieve flawless results. Finally, different versions

and upgrades of the current HB can be built. For instance, protecting the whole head by

means of a hat or dividing the HB in more/less pieces could represent new goals for future

studies.

23

6 Conclusion

The simulations show important results about the safety properties of the HB in terms of

protection of the head during a fall. Damages to the brain and skull can be effectively reduced

by wearing the HB in several daily activities i.e. while walking or cycling. Max. linear

acceleration, brain strain and risk of brain damages can be respectively reduced up to 74%,

52% and 64% depending on the impact locations and fall conditions. The pin structure and

HB response to impacts characterize a device that can provide protection to the head by

deforming and regaining its original shape. However, more in-depth studies should focus on

testing specific configurations and material properties to achieve even better values that

would maximize the efficacy of the HB.

25

7 References

[1] World Health Organization, ‘WHO Global Report on Falls Prevention in Older Age.’, Community Health (Bristol)., p. 53, 2007.

[2] ‘Rates of TBI-related Deaths by Age Group — United States, 2001–2010 | Concussion | Traumatic Brain Injury | CDC Injury Center’. [Online]. Available: https://www.cdc.gov/traumaticbraininjury/data/rates_deaths_byage.html. [Accessed: 28-Feb-2017].

[3] M. Faul, L. Xu, M. M. Wald, and V. G. Coronado, ‘Traumatic brain injury in the United States: emergency department visits, hospitalizations, and deaths’, Centers Dis. Control Prev. Natl. Cent. Inj. Prev. Control, pp. 891–904, 2010.

[4] C. Van Doorn, A. L. Gruber-Baldini, S. Zimmerman, J. R. Hebel, C. L. Port, M. Baumgarten, C. C. Quinn, G. Taler, C. May, and J. Magaziner, ‘Dementia as a risk factor for falls and fall injuries among nursing home residents’, J. Am. Geriatr. Soc., vol. 51, no. 9, pp. 1213–1218, 2003.

[5] R. D. Seidler, J. A. Bernard, T. B. Burutolu, B. W. Fling, M. T. Gordon, J. T. Gwin, Y. Kwak, and D. B. Lipps, ‘Motor control and aging: Links to age-related brain structural, functional, and biochemical effects’, Neurosci. Biobehav. Rev., vol. 34, no. 5, pp. 721–733, 2010.

[6] World Health Organization, ‘WHO | Dementia’. [Online]. Available: http://www.who.int/mediacentre/factsheets/fs362/en/. [Accessed: 08-Mar-2017].

[7] ‘AANS - Sports-related Head Injury’. [Online]. Available: http://www.aans.org/patient information/conditions and treatments/sports-related head injury.aspx. [Accessed: 02-Mar-2017].

[8] P. Kannus, H. Sievänen, M. Palvanen, T. Järvinen, and J. Parkkari, ‘Prevention of falls and consequent injuries in elderly people’, Lancet, vol. 366, no. 9500, pp. 1885–1893, 2005.

[9] J. Olivier and P. Creighton, ‘Bicycle injuries and helmet use: a systematic review and meta-analysis.’, Int. J. Epidemiol., p. 153, 2016.

[10] ‘Hövding 2.0 | | Hövding.com’. [Online]. Available: https://shop.hovding.com/. [Accessed: 08-Mar-2017].

[11] R. Anderson, G. Ponte, and L. Streeter, ‘Development of head protection for car occupants’, Road Transp. Res., vol. 12, no. 1, pp. 41–48, 2003.

[12] J. Okan, ‘Development of a fall-injury reducing flooring system in geriatric care with focus on improving the models used in the biomechanical simulations and evaluating the first test area’, 2015.

[13] M. Fahlstedt, P. Halldin, and S. Kleiven, ‘The protective effect of a helmet in three bicycle accidents — A finite element study’, vol. 91, pp. 135–143, 2016.

[14] Lstc.com., ‘LS-PrePost Online Documentation | Overview.’, 2017. [Online]. Available: http://www.lstc.com/lspp/content/overview.shtml. [Accessed: 10-May-2017].

[15] Lstc.com., ‘LS-DYNA | Livermore Software Technology Corp.’, 2016. [Online]. Available: http://www.lstc.com/lspp/content/overview.shtml. [Accessed: 10-May-

26

2017].

[16] T. U. A. Beskow, ‘Hip impact of the FE-model THUMS’, 2016.

[17] S. Kleiven, ‘Predictors for traumatic brain injuries are not evaluated through accident reconstructions.’, Stapp Car Crash J., vol. 51, pp. 81–114, 2007.

[18] P. Chan and E. Takhounts, ‘Development of a generalized linear skull fracture criterion’, no. January, 2007.

[19] R. A. Hospital, ‘Brain injury patterns in falls causing death’.

[20] M. Fahlstedt, K. Baeck, P. Halldin, V. S. J, J. Goffin, B. Depreitere, and S. Kleiven, ‘Influence of Impact Velocity and Angle in a Detailed Reconstruction of a Bicycle Accident’, pp. 787–799, 2012.

[21] H. J. Mertz, P. Prasad, and A. . Irwin, ‘Injury Risk Curves for Children and Adults in Frontal and Rear Collisions’, Proc. 41st Stapp Car Crash Conf., pp. 13–30, 1997.

[22] R. Schonnop, Y. Yang, F. Feldman, E. Robinson, M. Loughin, and S. N. Robinovitch, ‘Prevalence of and factors associated with head impact during falls in older adults in long-term care.’, CMAJ, vol. 185, no. 17, pp. E803-10, 2013.

27

Appendix A

Table of Contents

1 Falls-Related Head Injuries .............................................................. 31

1.1 Phenomenon and categories at risk ........................................................................... 31

2 Anatomy of the Head ...................................................................... 33

2.1 From the Hair to the Brain .......................................................................................... 33

2.2 The Brain ..................................................................................................................... 34

2.3 The Cerebrospinal Fluid (CSF) ................................................................................... 35

3 Biomechanics of the Head ............................................................... 37

3.1 Physics of Motion ........................................................................................................ 37

3.2 Dynamics of Impact .................................................................................................... 38

3.3 Brain Injury Types ..................................................................................................... 39

3.3.1 Translational Induced ........................................................................................... 39

3.3.2 Rotational Induced .............................................................................................. 40

3.4 Head Injury Criterion (HIC) and Parameters of Interest ........................................... 41

4 Prevention ..................................................................................... 43

4.1 Existing Solutions ...................................................................................................... 43

4.1.1 Helmets ................................................................................................................. 43

4.1.2 Protective Headbands ........................................................................................... 44

4.1.3 Shock Absorbing Material ..................................................................................... 45

5 Finite Element Method (FEM) ........................................................ 47

5.1 Human Head Model ................................................................................................... 47

6 References ..................................................................................... 49

29

List of Abbreviations

AIS Abbreviated Injury Scale

ATD Anthropometric Test Device

CNS Central Nervous System

CSF Cerebrospinal Fluid

DAI Diffuse Axonal Injury

EDH Epidural Hematoma

FE Finite Element

FEM Finite Element Method

HIC Head Injury Criterion

HIC15 Head Injury Criterion over 15 ms

HIC36 Head Injury Criterion over 36 ms

ICH Intracerebral Hematoma

KE Translational Kinetic Energy

KTH Kungliga Tekniska Högskolan

PMHS Post Mortem Human Subject

RKE Rotational Kinetic Energy

SDH Subdural Hematoma

SP Technical Research Institute of Sweden

TBI Traumatic Brain Injury

WHO World Health Organization

31

1 Falls-Related Head Injuries

1.1 Phenomenon and categories at risk

According to the World Health Organization (WHO), falls represent one of the most

prominent causes that lead to unintentional injury [1]. In particular, they are the leading

cause to traumatic brain injuries (TBI) [2]. Only in the USA, 595 095 cases of TBI due to falls

were registered between 2002 and 2006 [3]. Among the people who are majorly affected by

it, the elderly stand out as the category that is mostly at risk [1]. The graph in Fig.1 depicts a

clear situation that highlights this aspect. Specifically, Fig.1 addresses the US population in

2011 and shows the deaths related to unintentional injuries per 100 000 by age. Moreover,

TBI related deaths have also a major impact on the elderly. In fact, in Fig.2 it can be noticed

that the 65+ years old group is the most affected and the rate is increasing year by year [4].

The causes behind these trends depend on different risk factors that are, among the others,

age-related and wellness dependent. In particular, at a late stage of life, the neuro and motor

systems start degenerating and cognitive capacities decline. As demonstrated by Seidler et

al. [5], the structure, function and biochemistry of the brain change having a drastic impact

on the motor performance of the subject: gait and balance are compromised so that the risk

of falling increases. Comorbid conditions such as arthritis and osteoporosis are common

among the elderly and imply higher frailty and decreased range of motion [6]. Furthermore,

polypharmacy, which consists in the intake of four or more prescribed medications [6],

represents a serious threat to the wellbeing and life quality of the elderly. In particular, drugs

acting on the central nervous system (CNS) such as antidepressants and narcotic analgesics

used to relieve pain are associated with the increase of injurious falls [7]. In fact, the CNS,

Fig. 2 - TBI related deaths per 100 000 by age between 2001 and 2010 in the USA [4].

Fig. 12 - TBI related deaths per 100 000 by age in 2011 in the USA [4].

Fig. 1 - Unintentional injuries deaths per 100 000 by age in 2011 in the USA [5].

Fig. 13 - TBI related deaths per 100 000 by age in 2011 in the USA [4].Fig. 14 - Unintentional injuries deaths per

100 000 by age in 2011 in the USA [5].

32

which consists of the brain and the spinal cord, is where the information coming from the

body to the brain and vice versa is transferred and processed. Motor signals represent part

of this information and thus when the CNS is altered, also mobility functions can be

compromised. Additionally, after a fall, the fear of falling again triggers a mechanism that

limits the person in her daily activities and lowers her autonomy. Furthermore, the post-fall

syndrome leads the person to perceive herself as a burden for the society so that depression

and immobilization occur, deteriorating the person from a psychological and physical point

of view. This condition further increases the risk of falling again [8]. Based on statistical

data, it has emerged that the mentioned problem is shifting towards a considerable

increment of incidence due to the worldwide ageing population phenomenon. In 2006, the

60 years old group hit approximately 688 million and it is estimated to grow up to two

billions by 2050. Among the 65 years old, 28-35% fall each year whereas for the 70 years old

group the rate rises up to 32-42% [1].

Another category that is characterized by the high risk of fall and consequent head injury is

represented by people who suffer from a certain disorder or disease. One of the most

common disorders that increases the risk of sustaining an injurious fall is dementia [9].

Especially in the late stage, this syndrome implies difficulties in ambulation besides a severe

cognitive functions deterioration. WHO has estimated that the population affected by

dementia will reach 75.6 million in 2030 and it will be tripled by 2050 [10]. Patients affected

by epilepsy are also very likely to experience a head injury due to the high likelihood of a

seizure event [11]. Additionally, a remarkable phenomenon is represented by the onset of

epilepsy after a TBI. Experiencing a head-related injury represents the primary cause of

epilepsy [12] and it appears clear how TBI can be both cause and consequence, activating a

circle in which the subject is majorly exposed to injurious falls accidents.

Finally, cycling represents not only a popular recreation activity but also an everyday green

choice of commuting. Beside the environmental and health benefits, this activity implies

several risks. Among these, head injuries account for a large part and can result in fatal

outcomes [14, 15]. The American Association of Neurological Surgeons has estimated that in

2009 in the USA the recreational activity/sport associated with the highest number of head

injuries was cycling, accounting for 85 389 cases treated at the emergency room [15].

33

2 Anatomy of the Head

In order to fully understand the relevance of head injuries and their consequences, an

overview of the anatomy of the brain and skull is provided in the following subsections.

2.1 From the Hair to the Brain

The brain can be defined as the operative center of the human body and intellect. It is a very

fundamental and delicate organ and for these reasons it is encased in several coverings.

Starting from the most outer layer and proceeding towards the inner part, the anatomy of

the head is articulated as depicted in Fig. 3. The cutaneous layer (skin that bears the hair),

the subcutaneous connective-tissue layer, and the muscle and facial layer form the 5-7 mm

scalp [16]. The skull is the hardest part and it is composed by three layers of which the ones

in contact with the upper and lower structures are made of compact bone whereas the one

in-between, the diploë, is constituted by spongy bone. The epidural space connects the skull

to three membranes that are called meninges and whose goal is to provide not only

protection but also support to the brain.

From the outer to the inner one, they are [16] :

▪ dura mater, tough fibrous layer that serves as inner periosteum of the cranial bones.

The subdural space connects it to the following layer.

▪ arachnoid, a spider-web membrane situated in the subdural space. The subarachnoid

space anticipates the last membrane.

▪ pia mater, thin inner layer that adheres to the brain and is rich in blood vessels.

34

2.2 The Brain

Gyri and sulci are respectively ridges and depressions in the cerebral cortex and give the

brain its characteristic appearance. The brain can be divided in five compartments based on

its structure and functionalities: cerebrum, cerebellum, midbrain, pons and medulla

oblongata [16]. The brain is composed for seven-eighths by the cerebrum that is where the

higher functions such as thought and action are based. The cerebrum is partly divided

through the falx cerebri into two main regions that are called right and left hemispheres and

their surface is made of gray matter (nerve-cells bodies), the cerebral cortex. Underneath the

cerebral cortex, the white matter (neurons myelinated axons) lays and connects to the other

parts of the CNS. Within the white matter, agglomerates of gray matter called nuclei are

present. The hemispheres are divided by the longitudinal cerebral fissure, a deep fold, and

connected by the corpus callosum, a mass of white matter. The brain can further be divided

in four lobes as shown in Fig. 4 [16]:

▪ frontal, mainly associated with thinking, problem solving, emotions, behavior,

movements, and decision making.

▪ parietal, primarily related to spatial perception, spelling, and sensation.

▪ temporal, strictly connected to memory, language understanding, and hearing.

▪ occipital, principally associated with vision.

Fig. 3 - Main components of the head anatomy [17].

35

The cerebellum (or “little brain”) is also divided in two hemispheres which are connected by

a structure called vermis. This compartment is highly folded and is associated with fine

movement, posture, balance and motor learning abilities [16]. The midbrain, pons and

medulla oblongata form the brain stem. Among others, this part mostly accounts for the

regulation of breathing, heartbeat and blood pressure. It also represents the junction

between the upper parts and the spinal cord [16].

2.3 The Cerebrospinal Fluid (CSF)

This colorless fluid is produced from the arterial blood in the choroid plexuses of the

ventricles of the brain. It circulates around the brain and the spinal cord in the subarachnoid

space to fulfill its functions. It provides not only nutrients but also a cushioning mechanism

that protects the brain from mechanical shock. Moreover, the CSF is responsible for the

homeostasis and metabolism of the CNS. Finally, it returns to the venous system through

the arachnoid granulation villi and drains into the lymphatic vessels around the cranial

cavity (intracranial space that contains the brain and the meninges) and spinal cord [16].

Fig. 4 - Lobes of the brain [59].


Fig. 5 - Example of translational impact.

Fig. 17 - Example of translational impact.Fig. 18 - Lobes of the brain [59].


37

3 Biomechanics of the Head

This chapter is also supported by the description of the biomechanics involved in head

injury-related accidents and the criteria that are commonly in use in the literature to make

the best evaluation about the head damages that could occur after different fall scenarios.

3.1 Physics of Motion

After providing the basic anatomy concepts required to understand the different kinds of

head injury, some words must be spent on the physics of motion that characterize the

moment of impact. The injury process can be determined by two main types of motion:

translational (Fig.5) and rotational (Fig.6) [18].

The translation does not imply any rotation and it is usually known as linear motion. In this

case, the velocity that characterizes every point of the body does not change otherwise the

body would deform or rotate. For a pure translational motion, the kinematic parameters

taken into account are the displacement, x, the velocity, v, and the linear (translational)

acceleration, a. They are described by the following general relations:

The rotation of the body, instead, implies a change in its angular orientation. The

characteristic parameters for this type of motion are the angular displacement, θ, the angular

velocity, ω, and the angular acceleration, α. The relationships that describe them are:

The angular velocity is related to the velocity at two different points of the body, A and B,

and the distance between the two, r. By definition:

v(t) = dt

dx(t) a(t) =

dv(t)

dt

ω(t) = dt

dθ(t) α(t) =

dt

dω(t)

(VA - VB) ω =

r



Fig. 6 - Example of oblique impact that implies translational and rotational motion.Fig. 5 - Example of translational impact.


Fig. 6 - Example of oblique impact that implies translational and rotational motion.

38

As it will be better discussed in the following sections, the rotational motion in the impact

provides the major contribution in the most severe brain injuries.

3.2 Dynamics of Impact

When a force is applied, the body is subjected to an acceleration and the fundamental

relationship between force, F, acceleration, a, and mass of the body, m, is given by Newton’s

second law of motion: F=ma. Similarly, when the body undergoes angular acceleration, α, a

torque, T, is generated: T=Iα. I represents the moment of inertia [18]. A body with mass that

is moving at a certain velocity v is characterized by an amount of translational kinetic energy

(KE) that equals KE=1/2*mv2. If the body is subjected to a rotation, it possesses a

rotational kinetic energy (RKE) that equals RKE=1/2*Iω2 [18]. During an impact, the kinetic

energy can follow two paths. In the first case, it can be transferred i.e. varying the velocities

of the colliding objects. In the latter case, instead, the energy is converted into work i.e.

deforming the objects. The second case represents a major problem for the head because in

this energy absorptive process the brain is highly deformed, leading to head injury. When

hitting the ground, the head is subjected to a variation of the kinematic variables that

characterized it at the initial condition. Another fundamental aspect in the fall kinematics

regards the fact that the intracranial response is highly influenced by the impact direction

[19]. In particular, radial (normal), oblique and tangential impacts are characterized by

different kinematics as shown in Tab. 3.

Tab. 1 - Impact directions and descriptive kinematics.

Impact direction Kinematics

Translational Rotational

Radial x

Tangential x

Oblique x x

It has been demonstrated that radial impacts affect mostly the skull causing fractures

whereas the rotational contribution causes the relative motion of the brain inside the skull

so that higher strains and damages originate in the brain. This brain behavior is due to the

fact that its shear modulus is five/six orders smaller than its bulk modulus, making it

possible to consider the brain as a fluid when deformations arise [20]. This statement

underlies the high sensitivity that the brain undergoes when subjected to rotational loading

and thus shear.

By all means the physical characteristic of the individual such as shape, mass and stiffness

of the head determine the final outcome of the impact but the forces, energy, impact

direction, strains, stresses, velocities and accelerations etc. at the moment of impact and

after it are fundamental to describe, understand and determine brain injuries.

39

3.3 Brain Injury Types

From a clinical point of view, brain injuries can be classified in two main categories: focal

(local) injuries and diffuse injuries. The first category involves local damages in the brain

and it is visible by naked eye. Among these:

▪ Epidural hematoma (EDH)

▪ Subdural hematoma (SDH)

▪ Intracerebral hematoma (ICH)

▪ Contusions (coup and contrecoup)

The latter category instead, includes lesions causing spread damages such as global

disruption of the tissue and it is usually invisible. It comprehends:

▪ Edema

▪ Concussion

▪ Diffuse axonal injury (DAI)

The following paragraphs resume the main head injuries that are respectively primarily

induced by translation and rotational kinematics [20].

3.3.1 Translational Induced

Epidural hematoma (EDH)

The bone fragments generated after the skull fracture can lead the blood vessels to rupture

(see Fig. 7) so that accumulation of blood in the area between the skull and the dura mater

can occur [21].

Fig. 7 - EDH, SDH and ICH damaged areas [25].

Anterior Subdural hematoma

Epidural

hematoma

Intracerebral

hematoma

Posterior

40

Contusion (secondary)

Very common lesion that is characterized by areas affected by necrosis, pulping, infarction,

hemorrhage and edema. As shown in Fig. 8, contusions can be differentiated into two

categories depending on where the site of injury is located: coup, close to the impacted area,

and contrecoup, in a remote region from the point of impact [22]. Primary injuries differ

from secondary ones due to the fact that TBI occurs immediately after the initial trauma

whereas in the latter case the onset of the injury is an indirect result of the trauma [23].

Skull fracture

Break in the cranial bones due to the absorption of the impact energy in the skull. Damages

to it can generate fragments that could hurt the structures below it such as the meninges,

blood vessels and brain [21].

3.3.2 Rotational Induced

Concussion

Mild TBI that involves immediate loss of consciousness after the impact. It is usually

classified as a mild injury that is not life-threatening but can induce severe damages.

Diffuse Axonal Injury (DAI)

Many of the axons present in the hemispheres and in the white

matter are disrupted due to i.e. tearing (Fig. 9). The loss of

consciousness follows immediately the impact and can last for days

or weeks. Post-traumatic amnesia, severe memory loss and motor

deficits may be present. After one month, the chances of surviving

are halved [24].

A B

Fig. 8 - Contusions: A, coup; B, contrecoup [60].

Fig. 9 - DAI representation [61].

Fig. 54 - DAI representation [61].Fig. 8 - Contusions: A, coup; B, contrecoup [60].



41

Contusion

As in the translational induced case.

Subdural hematoma (SDH)

Rupture of the bridging veins that are present in the subdural space (see Fig. 7) and that

supply the brain. One common cause consists in the brain excessive rotation that leads the

blood vessels connecting the brain and the skull to be torn apart. The mortality rate is greater

than 30% [24].

Intracerebral hematoma (ICH)

It is caused by fast changes in the head acceleration/deceleration and consists in the

damaging of the neurons and glial cells (parenchyma) within the brain (see Fig. 7) [24].

Edema

Brain swelling due to an accumulation of fluid in the brain [25].

Depending on the level of severity of the injury, a categorical index has been developed from

an empirical basis: the Abbreviated Injury Scale (AIS). The range of assigned values falls

between 0 and 6 where:

0 = no injury

1 = minor injury

2 = moderate injury

3 = serious injury

4 = severe injury

5 = critical injury

6 = fatal injury

3.4 Head Injury Criterion (HIC) and Parameters of Interest

Among the several available head injury criteria used to evaluate the risk of injury and the

degree of acuteness, the Head Injury Criterion (HIC) is widely used to predict the risk of

sustaining head injury and it is defined as:

where t2 and t1 represent respectively the final and initial time (in seconds) of the interval in

which the HIC reaches the peak value and a is the linear acceleration. In conformity with the

42

common use, the maximum time duration lasts 15 or 36 ms, contact time intervals in which

the probability of the injury occurs [26]. For adults, HIC36 and HIC15 should not exceed

respectively 1000 and 700 for both mid-sized male and small female. This criterion is

suitable for only translational and frontal impacts and can be applied to predict skull

fractures and severe concussion [30, 31]. The threshold for mild concussion is set to HIC

values higher than 240 [28]. Apart from the HIC values mentioned above, it has been proven

that concussion is also well correlated to the first principal Green Lagrange strain [29].

Although its popularity, the HIC presents some limits. Among these, the fact that it is not

specific for a certain type of head injury makes it a topic for discussion in the today research

[30]. Furthermore, it only takes into account the linear acceleration whereas it has been

demonstrated that the majority of the injuries are related to rotational acceleration.

Among the others, some other parameters of interest that help to estimate the occurrence

and severity of the injury are the linear and rotational accelerations, von Mises stress, and

strain.

43

4 Prevention

Although many actions have been taken during the past years to decrease the risk of fall

injuries among the population especially within the elderly category, the statistics and data

show no major impact of the present solutions to solve the TBI issue.

Strength and balance training, use of walking aids, increment of vitamin D and calcium in

people’s diet, and the most of the age-friendly home/environment modifications can prevent

some fall hazards [43] but they do not provide any effective system to lower the entity of

head injury at the time of impact.

In the recent years, more age-friendly design has been implemented such as the energy

absorptive floors that will be treated in more detail in the following sections. Additionally,

the head injury severity issue can be addressed from another point of view by providing

protection directly to the head of the person. Helmets are a typical example of this kind of

approach and different types have been developed during the years. Some of these will be

discussed in the following paragraphs. The use of a helmet greatly reduces TBI [44];

however, the available solutions on the market give space for implementations from a design

point of view. The introduction of the innovative headband presented in this study could

provide a solution to the above-mentioned problems thanks to the adopted engineered

material that highly absorbs the energy of the impact.

4.1 Existing Solutions

4.1.1 Helmets

Different designs of helmets have been studied and tested by several researchers. Among

these, many sustain the importance of wearing a helmet to prevent and reduce head injuries

[14, 34, 48]. Although the effective protective action of wearing a helmet, many cyclists still

make no use of it. According to the study performed by Dagher et al. [46], the non-helmet

users among the patient admitted to the Montreal General Hospital between 2007-2011

belong to a younger and less educated category in which being single and unemployed were

also characteristic features. However, the reasons behind this kind of behavior are not clear

from the literature and can only be hypothesized. Design, functionality and efficiency are

believed to be the main contributing factors in the choice of a helmet.

44

Among the most recent and innovative solutions, Hövding, or better known as the airbag-

like helmet (see Fig. 10), represent an alternative to the common bike helmet. However, this

product presents different limits such as the high cost (2685 kr), the impossibility of re-using

it after it has been inflated, and the possibility of triggering it when the impact is not

happening.

4.1.2 Protective Headbands

Due to the severity of head injuries and their high occurrence, the attention of many

researchers and engineers has been drawn, especially in the last decades, to the

implementation of preventive devices. Headbands represent an attempt to prevent such

injuries and they have been developed for different occasions i.e. when playing soccer or

while sitting in the car [47]. An example of headband configuration is provided in Fig. 11.

This head-guard was invented and patented by Mary L. Aaron in 2002 [48] and it is mostly

suitable for children due to its light weight but it can be also worn by adults and the elderly.

However, no specific data about the testing of this device has been found in the literature.

Nonetheless, positive results have been observed about reducing brain damages among the

people who were wearing a protective headband during a car accident [47]. However, to the

best knowledge of the author, the present market does not offer a wide range of solutions

that could be addressed as satisfying in terms of efficiency and design for the targets and

occasions mentioned in the first section of this paper so it is strongly believed that there is

space for improvement from a design point of view.

Fig. 10 - Deflated and inflated configurations of the Hövding pre- and post- impact [63].

Fig. 63 - SmartCells solution for shock-energy absorbing floors. The flooring system with a detail of its cylindrical component is shown. When the load is applied, the cell deforms and regains its original conformation afterwards [64].Fig. 64 - Deflated and inflated configurations of the Hövding pre- and post- impact [63].



Fig. 11 – Headband design example; patent, USA, 2002 [48].Fig. 1068 - Deflated and inflated configurations of the Hövding pre- and post- impact [63].



Fig. 11 – Headband design example; patent, USA, 2002 [48].

Fig. 74 - SmartCells solution for shock-energy absorbing floors. The flooring system with a detail of its cylindrical component is shown. When the load is applied, the cell deforms and regains its

45

4.1.3 Shock Absorbing Material

The presented problem can also be addressed from another perspective: instead of creating

a product to be worn, engineers have developed a floor system capable of much higher

impact energy absorption, if compared with helmets, and able to provide good grip, stability

and comfort while walking on top [50]. Wright and Laing also show in their study a

considerable decrease in peak acceleration, peak force and HIC [51], making this system very

suitable for the presented scope.

An example is provided by the solution implemented by SmartCells, a company settled in

the USA and that has designed a floor that is characterized by a synthetic rubber material

and cylindrical components attached to the surface that exhibit a spring-like behavior when

loaded (see Fig. 12).

A similar and very promising flooring solution has been developed at KTH at the Neuronics

department by Prof. Svein Kleiven and Prof. Hans von Holst. It consists of rubber studs

aligned on multiple lines and they support the walking surface (see Fig. 13). When high loads

are applied i.e. during a fall, the pins are designed to bend making the floor softer and

allowing relative movement between the layers so that a higher impact energy absorption is

possible. After the load, the floor configuration goes back to its original shape [52]. The SP,

Technical Research Institute of Sweden, has recorded decrements equal to 60-75% and 39-

66% for respectively impact forces and accelerations [53]. The material and general design

of the KTH flooring system represents the key component in the innovative headband

proposed in this thesis work.

Fig. 12 - SmartCells solution for shock-energy absorbing floors. The flooring system with a detail of its cylindrical component is shown. When the load is applied, the cell deforms and regains its original conformation afterwards [64].

Fig. 93 - The black flooring system by KTH constitutes the middle layer between base and surface floor [53].Fig. 12 - SmartCells

solution for shock-energy absorbing floors. The flooring system with a detail of its cylindrical component is shown. When the

load is applied, the cell deforms and regains its original conformation afterwards [64].




Fig. 13 - The black flooring system by KTH constitutes the middle layer between base and surface floor [53].










47

5 Finite Element Method (FEM)

The Finite Element Method (FEM) represents a powerful numerical way to deal with

complex problems that are analytically hard to be solved. It allows to approach impact

scenarios in a cost-effective way and it also makes it possible to simulate a wide variety of

situations that present different degrees of complexity. If compared to the other two main

methods used in the field such as the Post Mortem Human Subject (PMHS) simulations and

Anthropometric Test Device (ATD) simulations (use of dummies), it can be stated that FEM

presents several advantages. Among these, it allows not only to simplify complex models

into simpler elements to obtain a solution for each of them, but it also provides the possibility

to virtually recreate dangerous scenarios without the necessity of using a living person or

animal. Additionally, ethical concerns and material body degradation problems related to

PMHS simulations can also be avoided [54]. FEM does not only achieve a good degree of

bio-fidelity in anatomy reconstructed models but it also allows to study the biological

response of the brain tissue through the brain injury predictors [55].

5.1 Human Head Model

The choice of using only the head as representative of the human subject in the impact

simulations finds its reasons in the fact that using a more complex human model is more

computationally and time expensive. Additionally, studies have showed that the use of the

upper part of the body such as the hands do not provide any protective benefit to the elderly

category that is addressed by this project. Motives behind this assertion rely for instance in

the delay of the physiological response of an old person during the fall. Lack of coordination,

damaged neurologic system, non-optimal muscle activation and tone also account as

explanations [56]. Additionally, as concern the cyclists as group of reference, several

previous studies such as by Fahlstedt [34, 49] have shown good results achieved in similar

simulations if compared to real life scenarios thus the choice of pursuing this method is

supported by the literature.

The model used for the head is the one developed at Kungliga Tekniska Högskola (KTH) by

Kleiven [29] and modified afterwards by Fahlstedt (Fig. 14) [57]. The correction regards the

skull compact and trabecular bone stresses that have been set to respectively 80 MPa and 32

MPa in order to take into account the possibility of loss in the capacity of load-bearing when

high contact loads are applied. Furthermore, the modelling of the scalp was modified by

using four solid elements made of Ogden material to simulate a more precise response [57].

The model represents an adult person head and it weighs about 4.56 kg. It comprises the

scalp, skull, meninges, bridging veins, brain, CSF and a simplified version of the neck that

includes the brain stem in the initial portion of the spinal cord. The model presents further

details such as the definition of parts that characterize the inner organs. For instance,

ventricles, thalamus, corpus callosum and contacts between the several parts have been

modeled. The model comprises 24 505 elements and 38 723 nodes. The scalp is of particular

interest as point of contact between the headband and the head. It is modeled using the

material 077 O-OGEN RUBBER and solid elements. However, for a more detailed

description of the model, the original mentioned articles should be considered.

48

A

B

Fig. 14 - Finite Element (FE) representation of the head model. A – the mesh of outer layer forming the scalp is clearly visible. B – Sagittal section of the head model. The inner organs and their mesh is visible.

49

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Appendix B

Stiffness curves – Baseline

Fig. 1 - Graph representing Curve 1 and Curve 2, load curves related to the rubber material that characterizes the pins and the solid layer in the HB. For confidentiality reasons, the Force values have been masked.

57

Appendix C

Fig. 1 – Translational acceleration [g] obtained from the KTH floor system in a crown impact.

0

20

40

60

80

100

120

140

0 0,002 0,004 0,006 0,008 0,01 0,012 0,014 0,016 0,018 0,02

Tran

slat

ion

al A

ccel

erat

ion

[g]

Time [s]

Translational Acceleration, Crown impact

Orig 11952

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