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Biomechanical Analysis of Stress Shielding and Optimal Orientation for
Short and Long Stem Hip Implants
by
Peter Goshulak
A thesis submitted in conformity with the requirements for the degree of
Master of Health Science in Clinical Engineering, Graduate Department of the
Institute of Biomaterials and Biomedical Engineering, University of Toronto
© Copyright by Peter Goshulak 2014
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Abstract
Thesis Title: Biomechanical Analysis of Stress Shielding and Optimal Orientation for Short and Long Stem
Hip Implants
Copyright: Peter Goshulak, Master of Health Science in Clinical Engineering, Institute of Biomaterials and
Biomedical Engineering, University of Toronto, 2014
Introduction: Short-stem hip implants are becoming increasingly common since they are presumed to
reduce stress-shielding, a probable cause of decreased bone density and fracture. This thesis examines
how implant stem length and orientation affect stress-shielding.
Methods: 12 artificial femurs were implanted with either a short-stem or long-stem implant. Three
control femurs were left intact. Femurs were subjected to axial loading. Data from six surface-mounted
strain gauges was used to validate a finite element model. A range of orientations were simulated.
Stress was compared between implants and across each implant’s range of orientations.
Results: The model fit experimental data well (intact: slope=0.898, R=0.943; short-stem: slope=0.731,
R=0.948; long-stem: slope=0.743, R=0.859). No orientation reduced stress-shielding (α=.05, p>.05).
Stress-shielding was significantly higher in long-stem implants (63%, p<.001) than short-stem implants
(29-39%).
Conclusions: Short-stem implants reduce stress-shielding compared to long-stem implants; orientation
has no effect. Short-stem implants are more likely to maintain long-term bone strength.
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Acknowledgements
I am incredibly grateful for the help and support of a number of extraordinary individuals. The work I’ve
completed over the past two years would not be remotely possible without the input of these people.
To my supervisor Dr. Emil Schemitsch: thank you for the opportunity to work on such an interesting,
impactful project. Your supervision has not only taught me about clinical orthopedics, but guided me in
what constitutes quality scientific research. It has been a true privilege working under your guidance.
To my committee members Dr. Jan Andrysek and Dr. Marcello Papini: thank you for the excellent
suggestions and insightful questions, from both meetings and emails. You have reinforced the
importance of thoroughness and due diligence in the research process – I could not be more grateful. An
additional thank-you goes to Dr. Cari Whyne, for providing suggestions for my committee members and
agreeing to act as external examiner.
To Dr. Radovan Zdero: your willingness to help seems to be endless. Thank you for the comprehensive
planning, making all the behind-the-scenes arrangements, and reviewing countless abstracts, drafts, and
posters. Every student should have an advisor of your dedication and caliber.
To Saeid Samiezadeh: your guidance in FE methods has propelled my abilities beyond what I thought
myself capable. Thank you for your patience and willingness to teach; I appreciated every minute.
To Dr. Mina Aziz: thank you for your incredible help in mechanical experimentation; you streamlined the
entire process from day one, and provided some great conversations. Best of luck in I.M. – you will
absolutely excel.
To Patrick Henry: thank you for your dedication in making sure I had everything I needed throughout the
testing process. You went beyond simply delivering tools, to teaching me about each one – thank you.
To Dr. Omar Dessouki and Dr. Mike Olsen: thank you for your expertise in selecting implants and
performing the procedures – your guidance and experience is extremely appreciated.
To my family and friends: I could not have made it through the last two years without your endless love
and support. Whether your (feigned?) interest in listening to the same “here’s why hips are neat” spiel,
your acceptance of “nothing is new, but I’m almost done my thesis” as a valid excuse for a whole year,
or simply delivering food to my door: you’ve kept me sane and focused. Thank you to everyone.
Lastly, to Marty: thank you for reminding me, no matter how heavy things get, that if you put your mind
to it, you can accomplish anything.
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Contents Abstract ......................................................................................................................................................... ii
Acknowledgements ...................................................................................................................................... iii
List of Figures .............................................................................................................................................. vii
List of Tables ................................................................................................................................................ ix
List of Terms and Abbreviations ................................................................................................................... x
1. Introduction .......................................................................................................................................... 1
1.1. Motivation ..................................................................................................................................... 1
1.1.1. Hip replacement surgery ....................................................................................................... 1
1.1.2. Stress Shielding ..................................................................................................................... 1
1.1.3. Smith & Nephew SMF Hip System ........................................................................................ 2
1.2. Goal of this thesis .......................................................................................................................... 3
1.2.1. Problem Statement ............................................................................................................... 4
1.2.2. Scope ..................................................................................................................................... 5
1.2.3. Clinical impact ....................................................................................................................... 5
1.2.4. Roadmap of thesis ................................................................................................................ 5
2. Background Information and Current Literature .................................................................................. 7
2.1. Normal Hip Anatomy .................................................................................................................... 7
2.1.1. Bony Anatomy ....................................................................................................................... 7
2.1.2. Muscular Anatomy .............................................................................................................. 10
2.2. Biomechanics of the Hip ............................................................................................................. 11
2.2.1. Supporting bodyweight ....................................................................................................... 12
2.2.2. Locomotion ......................................................................................................................... 14
2.3. Overview of the Hip Replacement Surgery Method ................................................................... 15
2.4. Modular neck options of the SMF Hip System ........................................................................... 16
2.5. Importance of alignment and positioning .................................................................................. 18
2.6. Testing methods.......................................................................................................................... 19
2.6.1. Live patient data collection ................................................................................................. 19
2.6.2. Mechanical testing .............................................................................................................. 20
2.6.3. Finite Element Analysis ....................................................................................................... 21
2.7. Summary of current knowledge gap ........................................................................................... 23
3. Method ............................................................................................................................................... 24
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3.1. Phase 1 Method – Mechanical Testing ....................................................................................... 24
3.1.1. Performing the implant procedures ................................................................................... 24
3.1.2. Strain gauge placement ...................................................................................................... 28
3.1.3. Axial loading experiments ................................................................................................... 30
3.2. Phase 2 Method – Finite Element Model Generation ................................................................ 32
3.2.1. Creating model geometry ................................................................................................... 32
3.2.2. Creating the experimental setup ........................................................................................ 33
3.2.3. Simulation parameters ........................................................................................................ 36
3.2.4. Method of load application ................................................................................................ 38
3.2.5. Simulation Outcomes .......................................................................................................... 38
3.3. Phase 3 Method – Analysis and Optimization ............................................................................ 39
3.3.1. FE model setup .................................................................................................................... 39
3.3.2. Design of Experiments ........................................................................................................ 40
3.3.3. Outcomes measured from DOE .......................................................................................... 43
3.3.4. Optimization........................................................................................................................ 44
4. Results ................................................................................................................................................. 47
4.1. Phase 1 Results – Mechanical Testing ........................................................................................ 47
4.2. Phase 2 Results – Finite Element Model Generation .................................................................. 47
4.3. Phase 3 Results – Analysis and Optimization .............................................................................. 50
4.3.1. Primary Analysis of Calcar Stress Shielding ......................................................................... 50
4.3.2. Optimal Implant Orientations ............................................................................................. 54
5. Discussion ............................................................................................................................................ 56
5.1. Summary of Main Findings ......................................................................................................... 56
5.2. Phases 1, 2 Discussion - Correlation of Model to Experiments .................................................. 56
5.3. Phase 3 Discussion - Analysis and Optimization ......................................................................... 57
5.3.1. Mean and standard deviation of stress .............................................................................. 57
5.3.2. Distal load transfer .............................................................................................................. 61
5.3.3. Overall peak stress .............................................................................................................. 62
5.3.4. Significance of Implant Orientation and Optimization ....................................................... 63
5.4. Clinical Implications and Future Directions ................................................................................ 64
5.5. Limitations and Sources of Error ................................................................................................. 68
5.5.1. Phase 1 – Mechanical Testing ............................................................................................. 68
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5.5.2. Phases 2, 3 – Finite Element Model .................................................................................... 72
6. Conclusion ........................................................................................................................................... 75
7. References .......................................................................................................................................... 76
8. Appendix ............................................................................................................................................. 85
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List of Figures
Figure 1.1: The SMF (left) and Synergy (right) Hip Systems.......................................................................... 4
Figure 2.1: Anterior view of proximal left femur .......................................................................................... 9
Figure 2.2: Superior view of left femur; proximal femur (hip) in green, distal femur (knee) in grey ........... 9
Figure 2.3: Axes of rotation of the hip joint ................................................................................................ 11
Figure 2.4: Synergy stem surgical technique [36] ....................................................................................... 16
Figure 2.5: The SMF implant (top left) and head-on views of the seven modular neck orientations [22] . 17
Figure 2.6: Effect of poor (lower) and adequate (upper) anteversion on increasing torque-arm length
[45] .............................................................................................................................................................. 19
Figure 3.1: Intact and SMF-implanted femurs, mounted in cement blocks ............................................... 25
Figure 3.2: Medial views of SMF-implanted femurs (necks removed), in neutral (top row) and anteverted
(bottom row) alignment ............................................................................................................................. 26
Figure 3.3: Overhead view of the four implant and neck combinations .................................................... 28
Figure 3.4: Left: strain gauge placement for the SMF femur; adapted from [22]; Right: image of
instrumented femur .................................................................................................................................... 29
Figure 3.5: Selection of four load conditions considered (simplified, lateral views); “Condition 4” was
ultimately selected ...................................................................................................................................... 31
Figure 3.6: Parameters used to osteotomize the FE femur model (anterior view) .................................... 34
Figure 3.7: Parameters used to position the FE implant stem (superomedial view) ................................. 35
Figure 3.8: Fully meshed model .................................................................................................................. 37
Figure 3.9: Calcar region (left); medial femur (middle); osteotomized femur (right) ................................ 40
Figure 3.10: DOE levels for the four implant cases; the Synergy DOE was simplified to a single point per
anteversion level ......................................................................................................................................... 42
Figure 3.11: Options used in the Minitab “Response Optimizer” function ................................................ 45
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Figure 4.1: Correlation plot for No Implant (control) specimens ............................................................... 48
Figure 4.2: Correlation plot for SMF-implanted specimens........................................................................ 48
Figure 4.3: Correlation plot for Synergy-implanted specimens .................................................................. 49
Figure 4.4: Boxplot of average percent change in mean calcar stress; all groups significantly different
(p<0.001) except SMF-A/SMF-R (p>0.05) .................................................................................................. 51
Figure 4.5: Boxplot of average percent change in standard deviation of calcar stress .............................. 52
Figure 4.6: Boxplot of neutral stance change in peak stress location along medial shaft; negative values
mean distal shift .......................................................................................................................................... 53
Figure 4.7: Interval plot (95% C.I.) of peak stress in remaining femur; 106 MPa is ultimate tensile
strength ....................................................................................................................................................... 54
Figure 4.8: Interval plot (95% C.I.) of percent change in mean calcar stress by implant anteversion ....... 55
Figure 5.1: Anterior view of calcar Von Mises stress (MPa) in (left – right) Intact, SMF-Neutral, SMF-
Anteverted, SMF-Retroverted, and Synergy implanted femurs ................................................................. 58
Figure 5.2: Medial view of calcar Von Mises stress (MPa) in (left – right) Intact, SMF-Neutral, SMF-
Anteverted, SMF-Retroverted, and Synergy implanted femurs ................................................................. 58
Figure 5.3: Medial view of peak Von Mises stress (MPa) in (left – right): Intact femur, SMF-implanted
femurs (SMF-N, SMF-A, SMF-R), and Synergy-implanted femur ................................................................ 61
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List of Tables
Table 4.1: Results of correlation plots ........................................................................................................ 49
Table 4.2: Results of DOE analysis; data is given as Mean (Standard Deviation) ....................................... 50
Table 4.3: Adjusted P-values for pairwise comparisons of mean calcar stress .......................................... 51
Table 4.4: Adjusted P-values for pairwise comparisons of standard deviation of calcar stress ................. 52
Table 4.5: Adjusted P-values for pairwise comparisons of standard deviation of calcar stress ................. 53
Table 4.6: Optimal implant angles to reduce calcar stress shielding .......................................................... 54
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List of Terms and Abbreviations
Abduction: rotation of a joint in the frontal plane, away from the midline; in the hip, lifting the leg outward
Adduction: rotation of a joint in the frontal plane, toward the midline; in the hip, moving the leg inward
Acetabulum: cup-like structure of the inferolateral pelvis; interfaces with the femoral head to form the hip joint
Anterior: towards the front of the body
Anteversion: rotation which moves the indicated structure anteriorly
BMD (Bone Mineral Density): measurement of the mineral content in bones; indicative of bone strength and integrity
Calcar: the proximal medial region of the femur, inferior to the femoral neck
Cancellous bone: soft, lightweight interior of bones
Cortical bone: hard, compact exterior of bones
Elasticity/elastic modulus: stiffness of a material, as stress applied per unit of strain: (stress)/(strain); units MPa, GPa
Extension: rotation of a joint to straighten or increase the angle between components; in the hip, posterior rotation in the sagittal plane, as in standing from a seated position
External Rotation: axial rotation of a joint away from the midline; in the hip, lateral rotation in the transverse plane, as in opening the toes outward
FE (Finite Element): computational method of modelling continuous 2D or 3D geometry as a finite number of simple elements, comprising a “mesh”
Femoral head: ball-like structure of the proximal medial femur; interfaces with the acetabulum to form the hip joint
Femoral neck: structure connecting the femoral head to the proximal medial femoral shaft
Femoral shaft: long bony structure between the proximal (hip) and distal (knee) ends of the femur
Flexion: rotation of a joint to decrease the angle between components; in the hip, anterior rotation in the sagittal plane, as in a kicking motion
Frontal Plane: vertical plane dividing the body into anterior and posterior parts
Greater Trochanter: superolateral aspect of the proximal femur; attachment point for muscles involved in hip abduction
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Inferior: towards the bottom of the body
Internal Rotation: axial rotation of a joint toward the midline; in the hip, medial rotation in the transverse plane, as in pointing the toes inward
Lateral: away from the midline
Lesser Trochanter: posteromedial projection of the proximal femur; attachment point for muscles involved in hip flexion
Medial: toward the midline
MPa (Mega Pascals): unit of measure for stress or pressure; equivalent: 106 Pa, 106 N/m2, 1 N/mm2; similarly, GPa (Giga Pascals): 109 Pa
Neutral (gait): orientation of the hip joint between flexion and extension, as in standing straight; (neck version): modular neck angle between anteversion and retroversion
Posterior: towards the back of the body
Retroversion: rotation which moves the indicated structure posteriorly
Sagittal Plane: vertical plane dividing the body into left and right parts
SMF (Short Modular Femoral): short stem hip implant with modular neck options
Strain: measure of change in length over original length: (Lfinal – Linitial)/Linitial ; dimensionless
Stress: measure of force per unit area: F/A ; units MPa
Stress Shielding: negative effect of implants where stress is absorbed by the implant rather than the existing bone; leads to decreased bone strength
Superior: towards the top of the body
Synergy: long stem hip implant
Transverse Plane: horizontal plane dividing the body into superior and inferior parts
UTS (Ultimate Tensile Stress): maximum tensile stress which can be withstood by a material before failure
Valgus: outward angling of a distal bone or joint, relative to normal alignment; in the knee, produces knock-kneed stance
Varus: inward angling of a distal bone or joint, relative to normal alignment; in the knee, produces bow-legged stance
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1. Introduction
1.1. Motivation
1.1.1. Hip replacement surgery
Hip replacement surgery is one of most common joint replacement procedures in the world, with the
total number of worldwide surgeries increasing by approximately 25% to 154 patients in a population of
100,000, and nearly doubling in the US to 184 in 100,000, from 2000-2009 [1] [2]. Surgeries are often
performed due to joint pain or arthritis, loss of joint function [3], bone fractures [4], or other conditions
such as osteonecrosis [5] [6]. It is estimated that 90% of patients who undergo hip replacement surgery
experience good results, with some implants lasting 20-25 years [3]. However, 18 out of 100 hip
replacements in the US must be revised at some point [7]. Loosening of the implant and post-operative
periprosthetic fracture are more common causes for revision than failure of the implant itself [7] [8] [9].
Post-operative periprosthetic fractures of the remaining femoral bone are prevalent in 1% of primary
hip replacements and 4% of revision procedures [10]. This failure is common among the elderly
population as bone quality is significantly deteriorated [4]. Hip replacement is also becoming more
prevalent in a much younger population than would historically have hip replacements [11] [12].
Patients younger than 30 years of age present a challenge in treating end-stage hip disease due to their
poor bone quality, though patients in childhood and adolescence have also received hip replacements
due to juvenile rheumatoid arthritis [11]. In order to facilitate the inevitable future revision surgeries in
such young patients, maximum bone stock must be conserved [13].
1.1.2. Stress Shielding
Since hip replacements are becoming more common in a younger population, there is a demand for hip
implants which allow for a more active lifestyle [14]. Maintaining bone strength has always been a
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concern of load-bearing prostheses, but an increasingly active patient population increases this demand
since high-stress activities are common such as running or sports; therefore, implant designs should aim
to maximize the load-bearing capabilities of the bone. A contrasting, but similar concern is that patients
today will typically weigh 20% more than patients several decades ago, therefore, overall bone strength
must be conserved [15].
Stress shielding is an adverse long-term effect which leads to decreased bone strength. The metal
prostheses have a much higher stiffness than bone, and therefore are less likely to deform under the
stress of load-bearing activities [16]. By absorbing the majority of the load, the implant reduces the
mechanical force transmitted to the bone itself. Since bone requires continual mechanical stimulation to
remodel and regrow, a stress-shielded implant site will gradually lose bone density, known as bone
resorption [17]; this phenomenon is especially prevalent in the medial proximal femur, known as the
calcar region [18] [19]. This decreased bone density may lead to aseptic loosening and stem migration,
as well as periprosthetic fractures [18] [20] [21]. Stress shielding can be reduced by increasing the load
transmitted to the bone; however, it is important not to increase the load excessively, as this may lead
to fracture. Therefore, the ideal stress distribution which implants should aim to replicate is the
physiologically natural stress distribution, or that of an unimplanted femur [16].
1.1.3. Smith & Nephew SMF Hip System
The Smith & Nephew Short Modular Femoral (SMF) Hip System (Smith & Nephew, Cordova, TN, USA
[22]) is a hip prosthesis which aims to address the concerns of bone conservation, as well as provide a
high degree of patient specificity by implementing a modular neck and head system. The implant length
is 20% shorter than conventional hip stems, and the surgical technique allows for a 10 mm higher neck
resection, resulting in greater bone conservation [13]. The SMF system is one of a number [6] [23] [24]
[25] of recent hip stems which uses a short, tapered stem, rather than a long, cylindrical stem. The SMF
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stem was chosen for this study since the thesis supervisor commonly uses the Smith & Nephew long
stem implants, and wished to compare two implants of the same brand. The short stem implant’s design
is intended to improve proximal stress distribution, thereby reducing the effects of stress shielding.
Since short stems are a relatively new concept, there are no studies demonstrating long term success of
this claim. However, short-term results are positive for other short stem systems: a two-year follow-up
study indicated a 0.4% incidence of both femoral fracture and revision surgery, compared to 3.1% as
seen in traditional long-stem total hip replacements [21]. Other reviews have shown short stem
implants to have high Harris Hip Scores of 93 [5] and 95 [26], compared to a score of 87 in long-stem
implants [27]; the Harris Hip Score is a clinician-measured outcome used to assess the results of hip
surgery, considering pain, function, absence of deformity, and range of motion [28]. The shorter stem
reduces intraoperative complications, since the implant is easier to insert, especially through a smaller
incision [24]. Additionally, the shorter stem reduces the amount of bone loss [21].
The concept of short stems, therefore, has a number of advantages. While the mechanism of stress
shielding is mostly understood as a whole, the exact degree to which short stems protect against stress
shielding has yet to be quantified in the general sense; additionally, the stress distribution created by
the SMF system itself has yet to be investigated at all.
1.2. Goal of this thesis
The goal of this thesis is to determine the stress shielding response of the SMF hip system. There are
currently no studies addressing the long term potential of this specific implant, and only a handful that
consider short stem implants. In order to determine whether the short stem design is an improvement
over a traditional long stem, this thesis will compare the stress shielding response of the SMF system to
an existing long-stem system: the Synergy hip system (Smith & Nephew, Cordova, TN, USA [29]; Figure
1.1). The ultimate outcome of this thesis is to recommend one implant over the other in terms of
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reducing stress shielding; a secondary outcome is to determine how implant orientation affects stress
shielding, and to recommend an optimal position.
Figure 1.1: The SMF (left) and Synergy (right) Hip Systems
1.2.1. Problem Statement
The question posed in this thesis is: “How do implant stem length and implant orientation contribute to
stress shielding in the calcar region?”
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1.2.2. Scope
This thesis is completed as part of a two year Masters of Health Science (M.H.Sc.) in Clinical Biomedical
Engineering degree, and is being completed in conjunction with a number of work terms. The scope is
therefore limited to a specific aspect of the SMF hip stem’s performance, as well as that of the Synergy
stem. Only the long-term effects of stress shielding are studied, using methods that have been
previously established and validated. Only artificial femurs and computer simulation are used; no live
specimens, cadavers, or patients were studied. While the conclusions drawn from the present research
aim to provide insight into stress shielding, the full impact of any recommendations can only be
observed through clinical implementation.
1.2.3. Clinical impact
The clinical implications of this research are a number of recommendations regarding implant selection
and orientation, as well as insight into how stress shielding is realized; as such, these results can be
immediately implemented by surgeons. While stress shielding as a phenomenon is mostly understood,
the methods used in this study provide detailed, comprehensive conclusions regarding a number of
implantation parameters relevant to practicing surgeons, specific to the SMF and Synergy hip systems.
Furthermore, the model developed in this thesis can be easily extended to study other parameters or
entire implant systems, helping further the clinical knowledge of stress shielding in hip implants as a
whole.
1.2.4. Roadmap of thesis
This thesis is divided into a number of chapters for ease of reading. Chapter 1 is an introduction to hip
implants and provides the rationale for this thesis. Chapter 2 reviews relevant anatomy, current
methods of studying hip implants, and provides some criticisms which this thesis aims to address.
Chapters 3 presents the three-phase research procedure, describing the methodology used in each step.
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Chapter 4 gives the results of each phase. Chapter 5 interprets the results of the research and provides
conclusions and recommendations as to how these results may be used clinically, as well as future
directions that may be pursued.
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2. Background Information and Current Literature
The following is a comprehensive review of the information required to complete the research involved
in this thesis, as well as an appraisal of existing studies and how this thesis differs.
2.1. Normal Hip Anatomy
The hip joint is a ball-and-socket joint which connects each lower limb to the pelvis. Specifically, it is the
articulation between the head of the femur and the acetabulum of the pelvis. The ball-and-socket
nature of the joint allows for rotation of the components in all three axes.
2.1.1. Bony Anatomy
The proximal articular surface of the hip joint is the acetabulum, part of the pelvic bone. The pelvic bone
is a fusion of three bones: the ilium, pubis, and ischium. The acetabulum is a roughly hemispherical cup,
with a lateral inferior opening. The articular surface of the acetabulum – the lunate surface – extends
through the anterior, superior, and posterior acetabular surfaces. The ball-and-socket nature of the joint
implies a congruency between the ball (femoral head) and socket (acetabulum); in reality, the unloaded
joints are non-congruent [30]. Once loaded, the acetabular geometry changes to maximize contact
surface area with the femoral head, reducing force concentrations. Additionally, this congruency during
loading provides stability to the joint, preventing translation perpendicular to the acetabular axis.
The distal articular surface of the hip joint is the femoral head, part of the proximal femur. The femur is
the longest, largest bone in the body. The proximal femur has several regions specifically relating to the
hip joint: the femoral head, femoral neck, greater trochanter, lesser trochanter, and femoral shaft
(Figure 2.1). The femoral head sits atop the femoral neck, which is angled in such a way to displace the
femur laterally from the pelvis and allow greater range of motion. This angle – the angle of inclination –
is measured in the frontal plane, and is typically 125° in normal, adult anatomy (Figure 2.1) [30] [31]. A
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large decrease in this angle results in varus hip alignment (distal femur angled inward), while a large
increase in this angle results in a valgus hip alignment (distal femur angled outward). This angle has a
significant effect on gait and the muscles necessary to maintain proper gait and stance. When viewed in
the transverse plane, the femoral neck also produces an angle with respect to the knee joint’s axis, or
the transcondylar axis. This angle – the angle of torsion – is typically 12-14° (Figure 2.2) [30]. The
resulting femoral head position is anteverted, or shifted anterior relative to the rest of the femur.
Significant variation in this angle results in “toe-in” (increased angle) or “toe-out” (decreased angle) gait,
in order to keep the femoral head within the acetabular cup [32].
Two types of bone material comprise anatomical bones: cortical and cancellous bone. Cortical bone is
the solid outer layer, and provides bones their strength and stiffness. Cancellous, or trabecular, bone is
the spongy material of bones, with lower density to reduce weight. In long bones such as the femur, the
cross-sectional thickness of cortical and cancellous bones varies along the bone length: through the
diaphysis (shaft), cortical bone is thicker to prevent bending, while in the epiphyses (proximal and distal
ends), cancellous bone is more prevalent to absorb compression [33].
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Figure 2.1: Anterior view of proximal left femur
Figure 2.2: Superior view of left femur; proximal femur (hip) in green, distal femur (knee) in grey
Angle of Inclination
Femoral Head
Femoral Neck
Greater Trochanter
Femoral Shaft
Lesser Trochanter (posterior)
Angle of Torsion
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2.1.2. Muscular Anatomy
There are 22 muscles associated with the hip joint, which act to provide stability and perform
movements of the joint. The ball-and-socket nature of the joint allows rotation of the femur in three
planes (Figure 2.3). Rotation in the sagittal plane is known as flexion and extension. Hip flexors such as
the psoas major, iliacus, and rectus femoris act to move the distal femur anteriorly, as in a kicking
motion. Hip extensors such as the gluteus maximus act to move the distal femur posteriorly, as in
standing from a seated position; hip flexors and extensors have significant roles in maintaining posture
and gait. Rotation in the frontal plane is known as abduction and adduction. Hip abductors such as the
gluteus medius and gluteus minimus act to move the distal femur laterally, raising the leg outwards; the
abductors play a significant role in maintaining posture. Hip adductors such as the adductor longus,
adductor brevis, and adductor magnus act to move the distal femur medially, closing the legs together.
Rotation in the transverse plane is known as internal and external rotation. External rotators such as the
piriformis, obturators, gemelli, and quadratus femoris act to rotate the femur laterally, as in pointing the
toes outwards. There are no specific internal hip rotator muscles, but rather internal rotation is a
secondary product of several other muscles acting together; this motion rotates the femur medially, as
in pointing the toes inwards. For the remainder of this thesis, the broad terms of muscle groups (hip
flexors, extensors, etc.) will be used instead of specific muscles. It should be noted that the actions of
many muscles depend on the orientation of the hip joint; for example, the gluteus medius and gluteus
minimus perform abduction when the hip is in extension, but perform internal rotation when the hip is
in flexion [34].
The range of motion of the hip joint is influenced by both the bony architecture and surrounding soft
tissues, including muscles, tendons, and ligaments. In general, passive range of motion is about 25%
greater than active range of motion, due to inefficient muscle contraction while near maximum
contraction [32]. The hip has a range of approximately 120° in flexion, with the knee fully flexed; this
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range is reduced when the knee is extended due to taut hamstring muscles. In extension, the hip has a
range of about 20°, though extension of the lumbosacral spine may provide some of this extension [32].
The hip has a 45° range of abduction, 30° range of adduction, and 45° range of both internal and
external rotation.
Figure 2.3: Axes of rotation of the hip joint
2.2. Biomechanics of the Hip
The hip joint has two major biomechanical functions: to support the body weight, and aid in locomotion
or other dynamic activities [31].
Flexion
Extension Abduction
Adduction
External Rotation Internal
Rotation
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2.2.1. Supporting bodyweight
Under static conditions, the hip joint typically supports the body in either single or double-leg stance;
most often, these stances occur with the joint in neutral alignment – that is, not in flexion or extension.
In double-leg stance, the body is supported mostly through passive means: rather than rely on muscle
contraction to maintain stance, the bony architecture and ligaments of the joint provide support,
minimizing the energy expenditure during double-leg stance [32]. This passive means of maintaining
stability is especially prevalent when considering stability in the frontal plane: the body’s center of mass
(COM) is located midway between the two hip joints, which cancels any force or moment imbalances
and maintains equilibrium. Losing balance in the frontal plane would require the body’s COM to shift
laterally beyond one of the legs, a significant margin.
Double-leg stability in the sagittal plane is also maintained passively, but is more easily perturbed. While
standing in good posture, the COM is located just posterior to the axis of hip rotation; this creates a
moment where the body tends to rotate posterior, causing hip extension. The iliofemoral ligament,
which passes anterior to the hip joint from above the acetabulum to below the femoral neck, stretches
passively to provide the counteracting anterior moment, thus maintaining balance. Hip flexor muscles
can also be actively recruited to maintain or regain normal posture if the COM is too far posterior or the
iliofemoral ligament is inadequate. Conversely, if the COM moves anterior to the hip’s axis of rotation,
hip extensor muscles must be actively recruited to regain stability.
Single-leg stability is a more complex task than double-leg stability, and requires more active muscle
recruitment to maintain. Stability in the sagittal plane is largely the same in single-leg stance as in
double-leg stance: the COM is typically located posterior to the hip’s sagittal axis of rotation, requiring
the iliofemoral ligament to provide the balancing anterior moment. In the frontal plane however, the
COM is no longer located between two supporting legs, but rather is medial to the single, supporting
13
leg. This creates a medial moment of rotation about the hip’s frontal axis, causing the pelvis to tend to
drop on the contralateral side. Stability is maintained by recruiting the hip abductor muscles of the
supporting leg, which acts to prevent the pelvis from dropping and “pull” the COM toward the ipsilateral
side; during static single-leg stance, the balancing leg is typically held in adduction to allow the upper
body to maintain an upright position. The “lever arm ratio” is the ratio of the perpendicular distance
between the COM and hip axis to the perpendicular distance between the abductor muscle and hip axis.
This ratio ranges from 1.25 to 3.00 [30] [32], meaning that the hip abductor muscles must act with 1.25-
3.00 times the force of the superincumbent body weight (the weight of the body supported by the hip,
not including the supporting leg – roughly 5/6 normal body weight [32]) to maintain equilibrium about
the hip axis.
As a result of these counteracting moments, a high compressive force is present between the femoral
head and acetabulum. In double-leg stance, forces range from 0.5 to 2.42 times body weight; in single-
leg stance, forces range from 2.0 to 5.4 times body weight [35]. More strenuous, high-impact activities
will have higher compressive forces: for example, jogging produces forces ranging from 4.33 to 5.84
times body weight. These compressive forces are distributed through the proximal femoral epiphysis by
a series of oriented trabeculae: the arcuate system, which resists tensile forces caused by bending about
the femoral neck; the medial system, which resists compressive forces in the femoral head; and the
lateral system, which resists compressive and tensile forces acting between the greater and lesser
trochanters [30]. Epiphyseal strength is highest where these systems intersect orthogonally, and
weakest where these systems are absent. One such region of absence is known as Ward’s Triangle,
located inferior to the femoral neck, which is especially prone to fracture in older patients with
decreased bone strength [30].
14
The medial offset of the femoral head in relation to the rest of the femur creates a bending moment
when the femur is vertically loaded. This can be visualized as the femur bending inwards at both ends
like a bow. This bending creates a compressive stress along the medial shaft, and a tensile stress
laterally [30]. The compressive stress is concentrated in the calcar region, while tensile stresses are high
near the greater trochanter due to the attachment of the hip abductor muscles. These and other forces,
depending on the activity and joint orientation, are manifested along the length of the femoral shaft;
other muscle and ligament attachments act to minimize adverse effects and stress concentrations.
2.2.2. Locomotion
The second major function of the hip is to aid locomotion, typically through walking or running. Due to
the alternation between single and double-legged support, gait can be thought of as “the process of
losing balance and regaining it” [32]. A full gait cycle for a single leg can be divided into stance phase,
approximately 60% of the cycle, and swing phase, approximately 40%. The stance phase can be further
subdivided into heel-strike, flat-foot, mid-stance, heel-off, and toe-off periods; the swing phase is less
clearly subdivided into acceleration, mid-swing, and deceleration periods [32]. Each hip joint goes
through a reciprocating cycle of flexion – neutral – extension – neutral – flexion; load is applied
throughout the range of flexion through extension by the supporting leg.
In addition to the range of flexion and extension in the sagittal plane, the hip undergoes rotation
fluctuations in the other two planes. The purpose of these fluctuations is to minimize displacement of
the COM, thus reducing total energy expenditure during gait [31]. In the frontal plane, the pelvis drops
on the swinging leg side during mid-stance: this reduces the potential vertical rise of the COM while the
body is directly above the leg; this is manifested as hip adduction in the supporting leg. Also in the
frontal plane, hip adduction during toe-off and swing deceleration of both hips keeps the knees and feet
closer to the midline: this reduces the side-to-side sway of the COM. In the transverse plane, internal
15
rotation of the supporting leg’s hip during toe-off effectively lengthens the leg: this reduces the vertical
drop in COM which occurs during full flexion and extension of the legs. Throughout the gait cycle,
internal and external rotation each have typical ranges of 4°, while abduction and adduction each have
typical ranges of 5°. The range of hip flexion is typically 25°, while extension is typically 10° [32].
2.3. Overview of the Hip Replacement Surgery Method
Hip stems are typically categorized as “cemented” or “uncemented”. Cemented stems use a fast-setting
cement as the bonding agent between implant and bone; uncemented stems have a porous outer
coating which bonds to the bone over time, and are “friction-fit” during insertion. The general surgical
method for an uncemented hip replacement surgery is given as follows; the images below correspond to
the Synergy stem, but the steps are similar for the SMF stem [36]. The surgeon exposes the hip using a
variety of potential approaches. After dislocating the joint, the acetabulum is reamed and an acetabular
cup component is impacted into the pelvis. The femur is then transected at the appropriate height and
angle for the selected implant (Figure 2.4 a). The femoral canal is then opened using a conical reamer
attached to a handle or electric drill (b). A series of successively larger toothed broaches are used to
shape the canal (c). The femoral stem is then inserted and impacted into the canal (d), and the
appropriate head is attached. In the case of a modular hip system, a trial neck and head are attached,
and the hip is reduced to check the implant’s function. Once an appropriate trial neck has been
determined, the actual neck and head are attached, and the hip is reduced and closed.
16
Figure 2.4: Synergy stem surgical technique [36]
2.4. Modular neck options of the SMF Hip System
In addition to providing a short, tapered stem design which differs from traditional long-stem implants,
the SMF allows for four different modular neck pieces to be combined in seven different orientations
(Figure 2.5). The “Standard” neck has a neck-shaft angle of 131°; the “High Offset” neck allows this angle
17
to increase or decrease by an additional 6°, resulting in an 8mm increase in either leg length (vertical) or
lateral offset (horizontal), depending on whether the neck is inserted right-side up or up-side down. The
“Right” and “Left” necks provide 10° of anteversion when inserted in the appropriate femur, or 10° of
retroversion when inserted in the opposite femur; additionally, they provide 6mm of lateral offset.
These two options can be reversed to provide equivalent ante- or retroversion while providing a 6mm
increase in leg length. It is important to note that the implant stem is inserted at approximately the
same angle of torsion (12-14° of anteversion) as the intact femur; thus, the additional anteversion or
retroversion provided by the modular neck options are in addition to this initial anteversion. Ultimately,
the surgeon has a wide variety of options to help restore normal, physiological anatomy in uneven hips.
This thesis will investigate a number of these options.
Figure 2.5: The SMF implant (top left) and head-on views of the seven modular neck orientations [22]
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2.5. Importance of alignment and positioning
The alignment and positioning of the prosthetic components in the hip joint is critical to maintaining
proper hip function. Proper positioning contributes significantly to the stability of the prosthetic joint,
which is especially critical early after an implant as the stem has not become fully bonded to the bone.
Dislocations are frequent occurrences, and increase in frequency with revision surgery [37] [38]. Joint
instability is a common complication in hip replacement surgeries, which contributes significantly to
dislocation [39]. 50%-70% of dislocations occur in the 5 weeks to 3 months postoperatively, suggesting
operative error in implant positioning [40] [41]. For example, it is difficult for surgeons to accurately
estimate femoral stem neck anteversion intraoperatively [42]. There is a mean surgeon estimation error
of 7.3° overestimation, which increases as anteversion decreases [43]. Another review found an 8.1°
difference between preoperative anatomical hip anteversion and postoperative anteversion [44]. These
discrepancies are significant as the optimal range for anteversion is 10-25° with respect to the
transcondylar axis [45]. Below 10° the joint will become progressively retroverted, which may cause
instability and lead to dislocation; anteversion over 25° is also associated with joint failure and
frequently requires revision surgery [45].
The effect of anteversion is readily demonstrated in the torque applied to the implant during stair
climbing. Figure 2.6 illustrates a transverse section looking down the right femur from the hip joint [45].
The lower implant neck represents little or no anteversion, and the upper implant neck represents
adequate anteversion (exaggerated). During stair climbing, the reaction force from the pelvis is applied
to the femoral head along the F vectors. A torque is therefore applied about the implant stem and
femoral canal. The magnitude of the torque is increased with the length of the torque arm. Insufficient
anteversion increases this torque arm length, thus increasing implant instability and loosening.
19
Figure 2.6: Effect of poor (lower) and adequate (upper) anteversion on increasing torque-arm length [45]
Femoral stem positioning is also critical to proper hip stability. A posterior approach by the surgeon has
been associated with better femoral stem positioning [46] [47]. However, the posterior approach is also
associated with a slightly higher incidence of dislocations [37]. Short stem prostheses have been shown
to have a wider varus-valgus range of alignment, as well as a greater overall femoral lateral offset [23];
this is due to the implant geometry, as the shorter stem allows greater varus-valgus rotation within the
medullary canal. Finally, there is a higher degree of difficulty in implanting uncemented stem
components, since the stem must be impacted into cancellous bone [48]. Therefore, the short-stem,
uncemented nature of the Smith & Nephew SMF Hip System requires increased care in positioning and
alignment of the stem component in order to reduce instability and the possibility of dislocation.
2.6. Testing methods
2.6.1. Live patient data collection
It is important to keep in mind that the goal of hip arthroplasty is to replace the normal function of a
human joint; therefore, results from actual patients should bear the most significance. Bone mineral
density (BMD) is a commonly used measure of stress shielding [18]; bone will remodel in response to
normal stress, though the precise mechanism is unclear [49] [50] [51]. BMD can be measured by dual-
20
energy X-ray absorptiometry, which can be combined with CT scans of the bone in question to create
patient-specific finite element models [20] [52]. One study of four long-stem implants showed
significant BMD reductions of 15-24% after two postoperative years in three of the four implants,
specifically in the calcar region [20]. While gathering data from live patients presents a realistic
environment for long-term investigation, the method has several disadvantages [53]: primarily, long-
term investigations will not yield rapid results, critical for developing new solutions; some patients may
opt out, require revision surgery, or die during the follow-up period; finally, significant differences in
patient demographics and anatomic geometry must be accounted for [54].
2.6.2. Mechanical testing
Mechanical testing of femurs can accurately demonstrate stress, stiffness, and strength patterns, and
aid in predicting biomechanical stability [54]. Mechanical testing typically employs axial compression or
torsion. Stiffness and strength in the specimen are recorded from the test machine, while strain gauges
mounted on the specimen surface are used to obtain strain measurements. However, mechanical
testing can be expensive and time consuming, and is subject to experimental error and material
variability in the specimens. Such variability is particularly prevalent in cadaveric testing, where both
material and geometric differences can significantly impact results [54]. To remedy this, synthetic
femurs can be used which present consistent material and geometric properties, as well as reduced cost
and ease of storage [55] [56]. Synthetic femurs are designed to match the biomechanical material and
geometric properties of live human specimens [57]. Mechanical testing on synthetic femurs has been
established as a valid method of gathering data on implanted femurs [54] [58].
The most common method of data collection for mechanical testing is performed with surface-mounted
strain gauges. Strain gauges are a simple, cost-effective method of measuring mechanical deformation
in mechanical specimens; they are applied to the surface using an adhesive and connected to a signal
21
processing device, which translates a change in electrical resistance into mechanical strain [59]. Strain
measurements have been used to measure the influence of short-stem implants on stress shielding. One
study using implanted cadavers [60] used six strain gauges located along the medial and lateral femoral
shaft; results indicated a reduction in stress shielding compared to long-stem implants, especially in the
proximal femur. However, the low number of specimens prevented precise values and significance from
being given. Another study using synthetic composite femurs [61] used eight strain gauges along the
medial and lateral femoral shaft; results again showed reduced stress shielding in the short stem
implant when compared to a long stem alternative, with the change in strain never exceeding 25% from
an intact femur except in the greater trochanter region. However, only one specimen was used for each
implant, so the authors noted that only qualitative data can therefore be collected. A significant
drawback of data collected solely from strain gauges is that they represent an average strain over a very
small area, thereby acting as a single point rather than an entire area.
2.6.3. Finite Element Analysis
While mechanical testing using strain gauges alone can provide strong conclusions regarding stress
shielding, computer models can be employed to simulate a wide variety of situations or make small
changes quickly and at little cost; this flexibility allows for a more comprehensive and definitive set of
results. Such simulations are termed Finite Element (FE) models. These models have been successfully
demonstrated as a valid simulation technique in determining stress and strain in the femur [54] [62]. FE
models are three-dimensional reconstructions of actual geometry, which can have external forces and
constraints applied to replicate actual loads. FE modelling of hip implants typically takes one of two
forms: reconstruction of patient-specific geometry using CT scans and BMD measurements, or using
synthetic or cadaveric specimens affixed with strain gauges.
22
Patient-specific FE models have been used to study the stress shielding characteristics of implants.
Results from cadaveric studies have validated the use of FE models to predict stress shielding behaviour,
and have confirmed the decreased proximal loading seen in previously discussed studies [63]. BMD
measurements can be used enhance FE models by indicating the material heterogeneity of the bone
[64]; this study demonstrated primarily distal-dominant load transfer in long stems, indicating the
presence of stress shielding. Results from several patient-specific analyses have been used to improve
implant design. One study used population-specific FE modelling to optimize implant geometry for the
population’s unique anatomy [65]; optimization focused on improving geometric fit to the anatomy to
improve implant stability, and improving proximal load distribution to reduce stress shielding. Patient-
derived FE models have also been used to study the impact of material selection on implant design [17];
while stress shielding can be reduced by decreasing the implant material’s stiffness, this will also cause
increased micromotion and implant loosening [17]. Therefore, implant design must take into account a
combination of material selection and geometry.
To provide more consistent results, synthetic femurs can be used in place of cadavers to develop FE
models. Synthetic femurs are advantageous since their inter-specimen variability can be 20-200 times
lower than cadaveric specimens [55]. They can be used to simulate multiple implant types and positions
for a single anatomy; however, implant design should not be based solely on a single femur geometry as
inter-population anatomical variations exist [63] [65]. Experimentally validated FE models of synthetic
femurs have been used extensively to study femoral fracture repair [62] [66] [67] [68]; however, this
methodology has yet to be used to study implants and their positioning within the femoral canal.
23
2.7. Summary of current knowledge gap
The current literature reveals that stress shielding is focused in the calcar region of the femur and leads
to decreased bone integrity. Stress shielding can be reduced by increasing the physiological load
transferred to the area. Implant orientation may have a significant role in determining both short-term
stability and long-term stress distribution via load transfer. In-vitro testing allows a wide range of results
to be obtained rapidly, and can be done by mechanical experimentation or FE modelling; short stem
implants have been studied by both of these methods, demonstrating improved stress shielding
characteristics compared to long stem implants. However, the two methods can be combined to yield
more precise measurements, as has been done to evaluate femoral fracture fixation techniques [66].
This thesis aims to address this knowledge gap by analyzing the SMF and Synergy hip systems to
investigate the precise degree of stress shielding caused by both implant stem length and overall
implant orientation.
24
3. Method
In order to measure the degree of stress shielding caused by implant selection and orientation, a three-
phase testing procedure was chosen which utilizes both mechanical testing and FE modelling. Phase 1
involves mechanical testing on synthetic femurs; the goal is to establish a collection of strain gauge data
for both implants, as well as an unimplanted control scenario, in a range of orientations and load
conditions. Phase 2 involves constructing an FE model of the experimental setup and using the strain
data to validate this model; the goal of the model is to replicate the mechanical experiments as closely
as possible. Phase 3 will then use this model to explore and analyze all possible implant scenarios; all
implants will be tested across their full range of physiologically-possible orientations, and these results
will reveal the exact effect the various parameters have on stress shielding. The conclusions drawn from
this analysis will determine the optimal implant selection and orientation.
3.1. Phase 1 Method – Mechanical Testing
The first step in creating a fully-validated FE model is determining the initial experimental response.
These results are the basis of an accurate simulation, and they should encompass a wide variety of
implant scenarios to provide the most thorough data. The nature of mechanical testing means there is
some inherent variability between test specimens; precautions such as high-precision measurements
were taken to minimize this error, but it could not be eliminated entirely. Therefore, multiple specimens
were tested for each scenario, and the results averaged.
3.1.1. Performing the implant procedures
Fifteen large left 4th generation synthetic femurs (Model #3406, Sawbones, Vashon, WA, USA) were
potted in cement blocks in 7° of adduction to simulate the single-leg stance of walking [60] [69]. The
femurs had a 16 mm diameter intramedullary canal, filled with solid cancellous foam of density 0.27
g/cm3; the cortical bone had density 1.64 g/cm3. The working length of the femurs was 330mm ±3mm
25
from cement block to the superior aspect of the greater trochanter (Figure 3.1). Three femurs were left
unimplanted to serve as control specimens. Six femurs were implanted with a size 5 SMF hip stem: three
in neutral alignment, three in maximum anteversion (Figure 3.2). Six femurs were implanted with a size
17 Synergy hip stem: similarly, three in neutral alignment, three in maximum anteversion. Implant sizes
were selected by preoperative templating by an orthopedic resident, and visually confirmed by a Smith
& Nephew representative. The implant procedures were performed by the primary author, a graduate
student, trained by an orthopedic resident experienced with both implant systems. Measurements and
photographs of the implants were confirmed as acceptable by an experienced orthopedic surgeon. It
can be seen in Figure 3.2 that the three stems seated in anteversion were not identically aligned. This
error would later be accounted for in the FE model.
Figure 3.1: Intact and SMF-implanted femurs, mounted in cement blocks
26
Figure 3.2: Medial views of SMF-implanted femurs (necks removed), in neutral (top row) and anteverted (bottom row) alignment
The manufacturer’s implant procedures were followed as closely as possible to maintain accurate
results. For the SMF stem, the primary indicator of success was a fully seated implant with full contact
along the inside lateral femoral cortex [22]; for the Synergy stem, a three-point fixation was indicated:
proximal-posterior, middle-anterior, and distal-posterior contact between the stem and femoral cortex
[36]. These contact conditions were confirmed visually by removing the implant, looking down the
femoral canal, and shining a flashlight through the cortical walls; the colour of transmitted light would
change from yellow, indicating cancellous bone, to purple, the colour of the cortical bone dye, in places
where the cancellous bone was completely removed and there was cortical contact. These conditions
were fully met for all implants in neutral alignment, but could only be met partially for those in
anteversion.
Some changes were made to the implant procedure, owing to the non-clinical conditions. The
osteotomy plane was extended to cut the greater trochanter. Normally this is an attachment point for
abductor muscles, but this muscle interaction was eliminated, simplifying the experimental load
27
condition and eliminating the need for this region; the greater trochanter was also not important for
implant stability as the implants were seated inferiorly. The orthopedic fellow performing implant
training noted that both cancellous and cortical bone were tougher than that of even a healthy patient,
and broaching the cancellous bone required a much greater effort; to reduce the amount of force
required, the canal reamer was used frequently to deepen the canal hole and clear debris. The femur
was also frequently overturned and shaken to clear debris. The excessive hammering caused the femurs
to loosen from the cement block; the femurs could not rotate since the cement was undamaged, but
most could translate about 1mm inferiorly out of the cement block. A steel plate was inserted below the
concrete block during loading to prevent this translation. The implications of these changes and
shortcomings to the procedure are addressed in Section 5.5.1.
Four implant “types” were selected for testing: three modular necks for the SMF system, and the
monolithic Synergy stem. The three modular necks chosen were “Standard”, “Left”, and “Right”. The
“Standard” modular neck produced no additional anteversion (henceforth this combination will be
called SMF-Neutral). When used in a left femur, the “Left” modular neck produces extra anterior offset,
or anteversion (SMF-Anteverted), while the “Right” modular neck produces posterior offset, or
retroversion (SMF-Retroverted). Figure 2.5 (above) shows an overview of all the various neck options,
and Figure 3.3 shows a summary of the four implant types selected in this study.
28
Figure 3.3: Overhead view of the four implant and neck combinations
Since it is rare that an implant is intentionally retroverted, not all combinations of modular necks and
implant angles were tested in the SMF-implanted femurs. In the neutral-alignment specimens (#1, 2, 3 in
Figure 3.2), only the neutral and anteverted necks were tested. In the anteverted specimens (#4, 5, 6 in
Figure 3.2), only the neutral and retroverted necks were tested. In total there were seven implant
groups: no implant, SMF-Neutral in neutral alignment, SMF-Neutral in anteversion, SMF-Anteverted in
neutral alignment, SMF-Retroverted in anteversion, Synergy in neutral alignment, and Synergy in
anteversion.
3.1.2. Strain gauge placement
Each femur was instrumented with five linear (Model #125UW, Vishay Precision Group, Malvern, PA,
USA; gauge length 3.18mm, width 4.57mm) and one rosette (Model #062UR; 45° rectangular, gauge
length 1.57mm, section width 1.57mm, overall width 10.67mm) surface-mounted 350 Ω strain gauges.
For the six femurs implanted with the SMF implant, three linear gauges were mounted along the lateral
cortex, located at 1.5, 3.5, and 6.0 inches (38.1, 88.9, and 152.4 mm) below the widest point of the
29
greater trochanter; these locations correspond to the middle, tip, and inferior to the SMF stem. Two
linear gauges were mounted on the medial cortex, located level with the two lower lateral gauges. The
rosette gauge was mounted in the calcar region, directly anterior to the lesser trochanter (Figure 3.4).
These locations were selected as areas of potential interest in strain measurement: the linear gauges at
the stem tip measured a predicted, and later confirmed, strain concentration, while the other linear
gauges were placed halfway from this tip to either stem top or cement block to measure the strain’s
dispersal over a distance while minimizing the effect of the cement block; the rosette gauge was place in
an area of complex geometry, since this was a predicted area of stress shielding.
Figure 3.4: Left: strain gauge placement for the SMF femur; adapted from [22]; Right: image of instrumented femur
Strain gauges were mounted in identical locations for the unimplanted control femurs. For the six
Synergy-implanted femurs, strain gauges were located in the same pattern but at different locations:
30
3.5, 7.5, and 9.5 inches (88.9, 190.5, and 241.3 mm) below the greater trochanter. Wire leads were
soldered to the strain gauge contacts, and connected via alligator clips to an 8-channel Cronos-PL data
acquisition unit (IMC MeβSysteme GmbH, Berlin, Germany). This unit was connected to a laptop running
IMC Device Control Software V2.6, which collected and stored the data for later analysis using FAMOS
V5.0 software (IMC MeβSysteme GmbH). Strain data was collected at 10 Hz, and stored in 30-second
segments.
3.1.3. Axial loading experiments
The specimens were secured distally to an angle vice mounted in an Instron 8874 mechanical tester
(Instron, Norwood, MA, USA). Three load angles were tested which represent a full gait cycle of single-
leg-stance loading: flexion (15°), neutral (0°), and extension (-15°); a higher angle for flexion, as is
physiologically typical of gait (Section 2.2.2), could not be achieved due to space constraints of the
mechanical testing platform. Quasistatic axial compression was applied by a flat steel block at subclinical
loads of 250N, 500N, 750N, and 1000N. A 50N preload was used to stabilize the load; the load then
ramped up over 10 seconds to the full loading level, held for 90 seconds, then ramped down to zero
over 10 seconds. The subclinical load level was chosen to prevent the femurs from undergoing plastic
deformation or fracturing, allowing them to be reused for the duration of the study. Multiple load levels
(250-1000N) were selected to confirm the linear elastic behaviour of the system. Each individual
experiment was repeated three times. Strain gauges were re-zeroed after each repetition.
The load was applied via a flat plate, rather than an acetabular cup substitute. Ideally, the flat plate
transmits load purely in the vertical direction, allowing lateral sliding; practically, some friction was
present, so petroleum jelly was applied to the plate to minimize sliding friction. Structurally, this
vertical-only load produces pure bending in the femoral shaft. In contrast, a cup-shaped load applicator
will produce horizontal reaction forces, with two negative consequences: buckling rather than bending
31
in the femoral shaft, and bending moment reactions in the load cell. While cup-style loading has been
used in previous studies [54] [69], other studies have acknowledged its drawbacks and used flat-plate
style loading instead [52] [60] [61]. Flat-plate style loading also had an unintended benefit of simplifying
the FE simulation and reducing computation time nearly ten-fold (Section 3.2.4). Figure 3.5 contains a
summary of various load application conditions considered.
Figure 3.5: Selection of four load conditions considered (simplified, lateral views); “Condition 4” was ultimately selected
32
3.2. Phase 2 Method – Finite Element Model Generation
Phase 2 involves creating an FE model of the experimental setup, then validating the model against the
experimental data obtained from Phase 1. Each experimental setup will be simulated individually, rather
than together with similar setups. This should compensate for the minute differences in experimental
setups inherent in the implantation process – for example, the variation in anteversion (Figure 3.2). A
number of setup parameters were defined, which fully describe the position and rotation of both the
synthetic femur and implant; each parameter can be individually adjusted based on measurements
taken from the experimental setup. Once the model has been validated, it can then be used in Phase 3
for a full analysis of the implant parameters of interest.
3.2.1. Creating model geometry
A previously-validated FE model of the Sawbones femur geometry was downloaded for use [54] [70]
[71]. The model is based on CT scans of a 3rd generation Sawbones femur; however, the geometry of the
4th generation is identical to that of the 3rd generation. The model contains solid bodies for both cortical
and cancellous bone.
FE models of both the SMF and Synergy stems were created by scanning the parts using a HandySCAN
3D handheld laser scanner (Creaform Inc., Levis, Quebec). The scanned files were imported into
Geomagic Studio 12 (Geomagic, Morrisville, NC, USA) where the model geometry was repaired. The
various SMF implant head and neck geometries were created in Solidworks 2013 (Dassault Systemes,
Waltham, MA, USA) using measurements and drawings available from the implant surgical technique
handbook [22]. The head and neck geometries were merged with the stem geometry in Geomagic,
creating a single part file for each of the four implants.
33
3.2.2. Creating the experimental setup
The femur model was imported into Ansys Workbench 15.0 (Ansys Inc., Canonsburg, PA, USA) and
positioned according to measurements from the experimental setups (Figure 3.6). A virtual origin was
defined at the most lateral surface of the greater trochanter; this origin was used as the basis for most
subsequent measurements. Two Ansys “design parameters” were used to define the femur’s
orientation: adduction and working length. Adduction defined the shaft angle, while working length
determined how much of the femur shaft is exposed distally. The femur was virtually osteotomized at
this working length, approximately 330mm from the top of the greater trochanter. This osteotomy
represented the solid concrete block from the experimental test. For the three control specimens, this
completed the model setup. For the implanted femurs, another virtual osteotomy was performed, this
time proximally. This osteotomy is identical to the cut done experimentally, and similar to that done in
the actual surgery. Two parameters defined this cut: osteotomy angle, which determines the angle of
the cut and varus-valgus (similar to adduction-abduction) rotation of the implant, and osteotomy level,
the distance from the virtual origin.
34
Figure 3.6: Parameters used to osteotomize the FE femur model (anterior view)
The appropriate implant model was then imported and positioned with respect to the osteotomized
femur (Figure 3.7). Three parameters were used to position the implant: anteversion, medial-lateral
position, and anterior-posterior position. Anteversion indicated the rotation about the femoral shaft to
produce an anterior offset by the implant head; the other two parameters are analogous to a Cartesian
coordinate system in the plane of the proximal osteotomy, allowing medial-lateral and anterior-
posterior positioning. Once positioned correctly, a “virtual surgery” was performed using a Boolean
subtract operation to remove cortical and cancellous bone from the regions occupied by the implant
stem.
35
Figure 3.7: Parameters used to position the FE implant stem (superomedial view)
Six surfaces were projected onto the cortical bone surface to match the locations of the six strain
gauges. Because Ansys calculated strain as the maximum – rather than average – nodal strain value for a
given surface, a small 1mm square surface was used instead of a patch equal in size to the strain gauge;
using a smaller virtual surface tended to give better approximations to the average strain over the larger
actual strain gauge. Each linear gauge was positioned at the same vertical distance along the femoral
shaft, while an anterior-posterior position parameter was used to correct for any errors in strain gauge
horizontal placement, improving overall accuracy to the experimental specimens. The rosette gauge had
both superior-inferior and medial-lateral position parameters. In total, seven parameters were used to
locate the virtual strain gauges. Surfaces corresponding to linear gauges measured normal strain in the
axial direction, while surfaces corresponding to the rosette gauge measured von Mises surface strain.
All 14 parameters (two to orient the femur, two for the osteotomy, three to orient the implant within
the femur, and seven to place the strain gauges) were measured for each of the 15 experimental
36
specimens, except osteotomy and implant orientation parameters for the three control specimens.
Working length was measured using Vernier calipers. All other measurements were taken from
photographs of the specimens, using MB-Ruler (Markus Bader, www.markus-bader.de). All
measurements were verified to ± 0.5° or ± 0.5 mm.
3.2.3. Simulation parameters
Material properties were assigned to the various solid bodies as follows [72] [73]: cortical bone Young’s
Modulus 16.7 GPa, Poisson ratio 0.26; cancellous bone Young’s Modulus 155 MPa, Poisson ratio 0.3;
implant stem, Titanium alloy ASTM F1472, Young’s Modulus 114 GPa, Poisson ratio 0.3; implant heads
and necks, Cobalt-Chromium-Molybdenum alloy ASTM F799, Young’s Modulus 240 GPa, Poisson ratio
0.3. All material properties were linear elastic and isotropic for simplicity, a common assumption when
modelling long bones [54] [71] [74]. Implant material properties were confirmed with a Smith & Nephew
engineer.
All contacts were designated as “bonded” type. This included cortical-cancellous, implant stem-cortical,
implant stem-cancellous, implant stem-implant neck. Initially, a non-linear frictional contact was
attempted between the implant stem and both bone types; however, excessively long solution times
and frequent program crashes resulted in this option being abandoned in lieu of the linear bonded
contact. The use of linear versus non-linear contacts is discussed in Section 5.5.2.
All bodies were meshed with tetrahedral 10-node elements (SOLID187) with maximum size 4mm (Figure
3.8). Meshes were patch-conforming. A mesh relevance of 80 was determined by iterative simulations
to produce a variation in cortical bone peak Von Mises strain of less than 1% in successive simulations.
The total number of elements ranged from 70,000 to 90,000.
37
Figure 3.8: Fully meshed model
The femur was fully constrained distally at the distal osteotomy surface, to simulate the concrete
loading block. A vector load was applied directly to the implant head, or to a similar-sized patch on the
superior surface of the femoral head in the intact control specimens. This load vector’s direction could
be varied to match the flexion, neutral, or extension stances that were used in the experimental
conditions. Four time intervals of loading were used, each corresponding to one of the 250N – 1000N
used experimentally. This entire setup is analogous to that described in Section 3.1.3.
38
3.2.4. Method of load application
As mentioned in Section 3.1.3 and Figure 3.5, a flat plate was used to apply the experimental load.
Initially in the FE simulations, a rectangular block was modelled to match this plate (steel, Young’s
Modulus 300 GPa, Poisson ratio 0.3, 4mm tetrahedral meshing). A non-linear frictionless contact was
applied between this block and the implant head, while frictionless-sliding (zero displacement) boundary
conditions were applied to the block sides to constrain it to vertical-only motion. The block was aligned
according to the flexion, neutral, or extension stance, and the load was applied to the top surface of this
block, compressing the implant head. The non-linear nature of this contact resulted in frequent non-
convergence of the solution, while successful solutions would typically run for one to two hours. The
flat-plate geometry could be simplified, however: since the flat plate produced no horizontal reaction
forces, only a downward force was ever applied. This force could instead be applied directly to the
implant head’s superior surface. While the physical model did present some sliding at the load interface,
resulting in a moving point of load application, the direction of loading remained constant; since the
virtual load was applied directly to the implant head surface, the direction of loading also remained
constant, regardless of any resultant sliding. This change in load application method eliminated the non-
linear contact condition, resulting in a fully linear model, with typical solution times of 5-10 minutes.
Slope and Pearson-R coefficient results between both non-linear and linear models were compared,
with no significant difference between the two.
3.2.5. Simulation Outcomes
Outcomes were acquired from the six surface patches corresponding to strain gauges; each virtual strain
gauge produced four surface strain results corresponding to each of the four load magnitudes, 250N –
1000N. The five “linear strain gauges” along the medial and lateral cortex measured normal strain in the
axial direction. The “rosette strain gauge” on the anterior calcar measured Von Mises equivalent strain.
39
These strain results were tabulated against the strain gauge data from Phase 1 experimental testing.
Correlation plots were created to determine the strength of the FE model.
3.3. Phase 3 Method – Analysis and Optimization
This final phase uses the validated FE model from Phase 2 to perform a full analysis of implant selection
and orientation. The end result of this phase is to accomplish the overall goal of this thesis by answering
the research question, as stated in Section 1.2.1: “how do implant stem length and implant orientation
contribute to stress shielding in the calcar region?”
3.3.1. FE model setup
For each of the five implant scenarios (no implant, SMF-Neutral, SMF-Anteverted, SMF-Retroverted,
Synergy), the same Ansys FE geometry was used as in Phase 2. The projected “strain gauge” surfaces
were removed, since Phase 3 analysis considered entire regions of the femur, rather than individual
points. Three such regions were defined: the calcar region, the medial femur, and the entire
osteotomized femur. The calcar region (Figure 3.9, left) is the primary region of interest of this study,
since prior studies have shown the highest stress shielding to occur in this region (Section 2.6). The
medial femur (Figure 3.9, middle) was chosen since preliminary simulations revealed intact femurs to
have a stress concentration in the calcar region, while implanted femurs showed a distal shift of this
stress concentration; monitoring stress in the medial femur allows this distal shift to be measured. The
osteotomized femur geometry (Figure 3.9, right) was considered the “overall” femur geometry for all
implant conditions, including the intact femur; this allowed stress to be compared in the same volume of
femoral bone. There was a 10 mm difference in osteotomy levels between SMF and Synergy implants,
but stress in the intact femur showed no difference between the two levels. The load was applied in an
identical manner as described in Section 3.2.4, except at a clinical-level load of 3000N, or 4x body weight
of a 75 kg patient. This load has been previously used since it represents a more realistic load [69] [75].
40
Figure 3.9: Calcar region (left); medial femur (middle); osteotomized femur (right)
3.3.2. Design of Experiments
A common tool used in gathering experimental data affected by multiple inputs is Design of Experiments
(DOE). Typically, two or more input parameters are analyzed with respect to one or more outputs. DOE
is commonly used when only a finite number of experiments can be produced, often due to time or
money. Each input is given a set of possible values, called levels. A full-factorial DOE will run an
experiment for each combination of levels across all inputs. Often, the number of experiments is
reduced by selecting certain combinations of levels which will provide sufficient information; however,
the low cost associated with running FE simulations allowed a full-factorial DOE to be performed (Figure
3.10).
41
In order to isolate the effects of implant orientation, several parameters were eliminated from the
analysis by holding them constant. These parameters were femoral adduction, femoral working length,
osteotomy angle, and osteotomy level (Figure 3.6, Figure 3.7). Additionally, medial-lateral implant
position was fixed for the SMF stem since the Smith & Nephew guidelines mandate that the lateral stem
should maintain full contact with the lateral cortex [22]. Since only femoral adduction and working
length were considered in the “No Implant” control case, a DOE was not necessary for this scenario, and
a single experiment was run to determine baseline stress results.
The remaining parameters were assigned levels which produce physiologically-possible implant
orientations. A physiologically-possible orientation is one where the implant does not excessively
impinge on the cortical bone, or protrude through the outer cortical surface. The range for anteversion
was set as 0°-15°, relative to the femur’s existing angle of torsion – in the Sawbones femur geometry,
this angle is 12°, making the total anteversion 12°-27° with respect to the knee’s transcondylar axis. The
remaining parameter level ranges were found by trial-and-error, and varied based on anteversion. Once
the allowable range was established for each parameter, five levels were evenly distributed through this
range. Since two inputs (anteversion, anterior-posterior position) were tested at five levels each, the
total number of experiments for each SMF stem was 25.
For the SMF-Anteverted and SMF-Retroverted implants, parameter ranges were chosen such that both
the implant stem and the resulting head position were within the 0°-15° anteversion envelope. While
the neck geometry is angled at ±10° when measured from the neck insertion point, the actual resulting
angle is ±6° when measured from the lateral aspect of the stem – this is where stem anteversion is
measured from. Therefore, the allowable range for SMF-Anteverted stems is 0°-9°, while the range for
SMF-Retroverted stems is 6°-15°.
42
The increased width of the Synergy implant resulted in a smaller allowable range of anterior-posterior
positions. This range was 1mm or less (Figure 3.10); therefore, the DOE was simplified from a two-input
design to a single-input design, measuring only anteversion. A single anterior-posterior position was
calculated based on anteversion, rather than varying between a maximum and minimum. Therefore,
only five experiments were run.
Figure 3.10: DOE levels for the four implant cases; the Synergy DOE was simplified to a single point per anteversion level
-4
-2
0
2
4
6
0 5 10 15
An
t-P
ost
po
siti
on
[m
m]
Anteversion [degrees]
Design of ExperimentsSMF - Neutral Neck
AP_min AP_max DOE points
-4
-2
0
2
4
6
0 5 10 15
An
t-P
ost
po
siti
on
[m
m]
Anteversion [degrees]
Design of ExperimentsSMF - Anteverted Neck
AP_min AP_max DOE points
-4
-2
0
2
4
6
0 5 10 15
An
t-P
ost
po
siti
on
[m
m]
Anteversion [degrees]
Design of ExperimentsSMF - Retroverted Neck
AP_min AP_max DOE points
-4
-2
0
2
4
6
0 5 10 15
An
t-P
ost
po
siti
on
[m
m]
Anteversion [degrees]
Design of ExperimentsSynergy
AP_min AP_max DOE points
43
3.3.3. Outcomes measured from DOE
As previously discussed, stress shielding represents a decrease in load transmitted to the region in
question; with hip replacement surgery, this phenomenon is most commonly seen in the calcar region.
The degree of stress shielding caused by an implant can be measured as a change in stress magnitude,
with respect to an intact femur [16] [61] [65]. While other studies have measured stress, or strain, at
predetermined points [60] [61] [63], the overall peak stress [65], or the qualitative location of stress
concentrations [64], the nature of FE simulation allows some more informative, comprehensive results
to be measured. With one exception, all results are measured in terms of stress or change in stress,
rather than strain. Strain is easier to measure in laboratory environments since data is collected via
strain gauges. However, fracture strength is given in terms of stress, and stress can be more easily
understood than strain [76]. Additionally, the true elastic properties of bone are not homogenous,
resulting in a non-linear relationship between stress and strain; stress parameters have been shown to
be better predictors of bone remodelling than strain parameters [52]. The non-stress result is a change
in length, given in millimetres. The regions indicated are illustrated in Figure 3.9. The results measured
from the DOE are:
Mean calcar stress: an average of equivalent Von Mises stress in calcar region nodes, which
indicates the overall amount of load transmitted to this region. Given as a percent change from
mean calcar stress in a non-implanted femur
Standard deviation of calcar stress: standard deviation of equivalent Von Mises stress in calcar
region nodes, which roughly indicates the variation of stress magnitudes in the region. This
indicates the presence, or absence, of stress concentrations. Given as a percent change from
standard deviation calcar stress in a non-implanted femur
Peak stress location: the quantitative indicator of proximal or distal shift in load distribution in
the medial femoral region. Given in mm.
44
Peak overall stress: the maximum stress in the entire osteotomized femoral cortical bone, which
indicates stress concentrations and sites of local bone failure. Given in MPa.
Mean and standard deviation of stress are calculated as changes in stress from an intact femur. This is
calculated as (“implanted femur” – “intact femur”) / “intact femur”. The mean, standard deviation, and
peak stress results were measured in all three stances (flexion, neutral, extension). The peak stress
location result was only measured in neutral stance; flexion and extension stances produced strong
torsion through the femoral shaft, and the peak stress location was always located distally, at the virtual
concrete fixation. These results were meaningless, and so were ignored.
These four results were collected from all DOE experiments for each of the four implant cases: 25
experiments for each of three SMF cases, five experiments for the Synergy case. Custom scripts were
written in the Ansys coding environment to calculate the results (Appendix). These results were
combined, creating four aggregate results for each implant case. These aggregate results were used as
the primary analysis of stress shielding, both in analyzing various orientations and comparing entire
implants. Analysis was performed in Minitab 16 (Minitab Inc., State College, PA, USA), using ANOVA at a
95% confidence interval. Pairwise comparisons were calculated using a General Linear Model, and p-
values were calculated using the Tukey test; statistical significance was set at p = 0.05.
3.3.4. Optimization
The results from the primary analysis were then used to perform optimization routines. The goal of this
part of the study is to determine the optimal orientation for each implant. This single orientation should
minimize stress shielding across all stance phases, without exceeding the fracture strength of cortical
bone. Orientation parameters were the same as in the primary analysis: anteversion and anterior-
posterior position for the three SMF implants, anteversion alone for the Synergy implant.
45
The optimization routine was performed in Minitab. The data was first defined as a set of DOE inputs
and outputs, then the “result optimizer” function was used. Three of the four result categories were
selected for optimization: mean calcar stress, standard deviation of calcar stress, and peak overall stress.
Peak stress location was not included because the variation between results was small and not
significant, with respect to orientation parameters. A screenshot of the Minitab “Response Optimizer”
window, showing the inputs and options used, is shown in Figure 3.11.
Figure 3.11: Options used in the Minitab “Response Optimizer” function
Nine total inputs were used: three for flexion (FLX) results, three for neutral (NEU), and three for
extension (EXT). All three stances were considered together since the goal of optimization is to find a
single orientation, for each implant, that optimizes results across all stances.
46
The first two inputs for each stance, “CalcarMean” and “CalcarSD,” refer to the mean and standard
deviation of stress in the calcar region. They are expressed as a percent change from the “no implant”
case, therefore the optimal value is 0% change. Since values can be above or below this value, the
“target” option was chosen. Upper and lower bounds of ±100% were chosen since all results fell within
this range.
The third input for each stance, “OvrPk,” refers to the overall peak stress for the entire femur. It is
expressed in MPa so it can be compared directly to the ultimate tensile strength of the synthetic bone
used in this experiment, 106 MPa [72]. Since all values below this can be considered safe, the
“minimize” option was chosen. The upper bound of 500 MPa was selected since most values fell within
this range, with a few outliers. A weight value of 10 was given to place more importance on values
closer to the target (Figure 3.11, bottom left).
47
4. Results
4.1. Phase 1 Results – Mechanical Testing
Data from the strain gauge channels were collected and recorded in IMC FAMOS software. The 30-
second interval of recorded data towards the end of the 90-second load application was selected for
analysis. Each channel was visually inspected to screen any abnormalities and to ensure that the 10-
second ramp-down in load was not selected. Data was exported into Microsoft Excel for analysis. Since
strain gauge placement was different from other studies, a direct comparison of surface strains is not
possible. Results from this mechanical testing can only be used to validate the FE model in Phase 2.
4.2. Phase 2 Results – Finite Element Model Generation
Strain results from the FE model were collected and plotted against the strain results from mechanical
testing. Each FE model data point was plotted against the corresponding experimental data point,
averaged across three experimental runs. There were seven implant groups tested: no implant, SMF-
Neutral in neutral alignment, SMF-Neutral in anteversion, SMF-Anteverted in neutral alignment, SMF-
Retroverted in anteversion, Synergy in neutral alignment, and Synergy in anteversion. Each implant
group had three specimens, each tested in flexion, neutral, and extension at 250N, 500N, 750N, and
1000N. Each experimental run produced six strain values, corresponding to the five linear and one
rosette strain gauges. Therefore, each of the seven implant groups produced 216 strain data points.
Results from the four SMF implant groups were grouped together under a single plot; results from the
two Synergy groups were also grouped together. Correlation plots for these three implants (no implant,
SMF, Synergy) are given in Figure 4.1 through Figure 4.3. A summary of the results are given in Table 4.1.
48
Figure 4.1: Correlation plot for No Implant (control) specimens
Figure 4.2: Correlation plot for SMF-implanted specimens
y = 0.8985x + 9.2015
-1500
-1000
-500
0
500
1000
1500
-1500 -1000 -500 0 500 1000 1500
FE S
imu
lati
on
[m
icro
stra
ins]
Experimental [microstrains]
Correlation: No Implant (n=3 specimens)
y = 0.7305x + 18.321
-1500
-1000
-500
0
500
1000
1500
-1500 -1000 -500 0 500 1000 1500
FE S
imu
lati
on
[mic
rost
rain
s]
Experimental [microstrains]
Correlation: SMF (n=12 specimens)
49
Figure 4.3: Correlation plot for Synergy-implanted specimens
Table 4.1: Results of correlation plots
No Implant SMF Synergy
Slope 0.898 0.731 0.743
Intercept 9.201 18.321 38.926
Pearson R coefficient 0.943 0.948 0.859
Slope Correction Factor 1.113 1.369 1.345
Data points 216 858 428
Outliers 0 6 4
Since the correlation plot slopes were not the ideal value of 1.0, a “slope correction factor” was
calculated as the inverse of the slope for each implant’s correlation plot. This was used to scale up all
future results to unity with experimental results. By using 4 load levels (250N – 1000N), the linearity of
the model was confirmed, allowing a single correction factor for all load levels. The “outliers” row
indicates the number of data points above a threshold of 1500 microstrains – typically these outliers
were two or three orders of magnitude greater than the rest of the results. These outliers skewed the
results heavily, and were removed from calculation. Overall, the Pearson R correlation coefficients of all
three plots showed strong correlation, with No Implant (0.943) and SMF (0.948) cases showing excellent
y = 0.7434x + 38.926
-1500
-1000
-500
0
500
1000
1500
-1000 -500 0 500 1000
FE S
imu
lati
on
[mic
rost
rain
s]
Experimental [microstrains]
Correlation: Synergy (n=6 specimens)
50
correlation. Therefore, the model is shown to be a valid predictor of experimental results, and can be
used for further analysis.
4.3. Phase 3 Results – Analysis and Optimization
4.3.1. Primary Analysis of Calcar Stress Shielding
Results from the DOE analysis are given below (Figure 4.4 through Figure 4.7). Table 4.2 contains a
summary of the Mean, Standard Deviation, and Peak Stress Location results. Overall peak stress will be
discussed qualitatively in Section 5.3.3. Boxplots were generated using one-way ANOVA in Minitab,
while the Peak Stress plot required an Interval Plot. Below each boxplot is a table with adjusted p-values
of pairwise comparisons. P-values for pairwise comparisons were calculated with the Tukey test at 95%
confidence interval. Significantly different groups (p < 0.05) are highlighted. Post-hoc two-tailed power
analysis of each pairwise comparison yielded 100% statistical power for all significantly different pairs.
Table 4.2: Results of DOE analysis; data is given as Mean (Standard Deviation)
No
Implant
SMF-
Neutral
SMF-
Anteverted
SMF-
Retroverted Synergy
Change in Mean Calcar
Stress (Figure 4.4)
0 % -38.97 %
(2.38 %)
-31.44 %
(2.48 %)
-29.87 %
(3.18 %)
-63.31 %
(1.91 %)
Change in Std. Dev. Calcar
Stress (Figure 4.5)
0 % -24.31 %
(6.70 %)
-14.18 %
(6.18 %)
-13.76 %
(7.81 %)
-43.11 %
(4.78 %)
Change in Peak Stress
Location (Figure 4.6)
0 mm -38.24 mm
(18.41 mm)
-46.59 mm
(5.38 mm)
-45.49 mm
(16.69 mm)
-140.67 mm
(9.63 mm)
51
4 Synergy3 SMF_R2 SMF_A1 SMF_N0 No Implant
0
-10
-20
-30
-40
-50
-60
-70
Implant
Av
g %Δ
Me
an
Ca
lca
r S
tre
ss
Figure 4.4: Boxplot of average percent change in mean calcar stress; all groups significantly different (p<0.001) except SMF-A/SMF-R (p>0.05)
Table 4.3: Adjusted P-values for pairwise comparisons of mean calcar stress
No Implant SMF-N SMF-A SMF-R
SMF-N < 0.001
SMF-A < 0.001 < 0.001
SMF-R < 0.001 < 0.001 0.238
Synergy < 0.001 < 0.001 < 0.001 < 0.001
52
4 Synergy3 SMF_R2 SMF_A1 SMF_N0 No Implant
10
0
-10
-20
-30
-40
-50
Implant
Av
g %Δ
Std
De
v C
alc
ar
Str
ess
Figure 4.5: Boxplot of average percent change in standard deviation of calcar stress
Table 4.4: Adjusted P-values for pairwise comparisons of standard deviation of calcar stress
No Implant SMF-N SMF-A SMF-R
SMF-N 0.007
SMF-A 0.260 < 0.001
SMF-R 0.289 < 0.001 0.999
Synergy < 0.001 < 0.001 < 0.001 < 0.001
53
4 Synergy3 SMF_R2 SMF_A1 SMF_N0 No Implant
50
0
-50
-100
-150
Implant
Ne
u.S
t. Δ
Pe
ak S
tre
ss L
oca
tio
n [
mm
]
Figure 4.6: Boxplot of neutral stance change in peak stress location along medial shaft; negative values mean distal shift
Table 4.5: Adjusted P-values for pairwise comparisons of standard deviation of calcar stress
No Implant SMF-N SMF-A SMF-R
SMF-N 0.082
SMF-A 0.019 0.256
SMF-R 0.023 0.396 0.999
Synergy < 0.001 < 0.001 < 0.001 < 0.001
54
Stance
Implant
3 Ex
tens
ion
2 Ne
utral
1 Flex
ion
4 Sy
nerg
y
3 SM
F_R
2 SM
F_A
1 SM
F_N
0 No
Implan
t
4 Sy
nergy
3 SM
F_R
2 SM
F_A
1 SM
F_N
0 No
Implan
t
4 Sy
nergy
3 SM
F_R
2 SM
F_A
1 SM
F_N
0 No
Implan
t
250
200
150
100
50
0
Pe
ak S
tre
ss [
MP
a]
106
Figure 4.7: Interval plot (95% C.I.) of peak stress in remaining femur; 106 MPa is ultimate tensile strength
4.3.2. Optimal Implant Orientations
Results from the “Response Optimizer” function are given in Table 4.6. Only the angle of anteversion is
reported since the anterior-posterior position is unlikely to be considered clinically, due to the small
range of possible positions and differences in patient anatomic morphology. Both the stem angle and
resulting head angle are reported; as stated in Section 3.3.2, the anteverted and retroverted necks
produce an additional ±6° of anteversion, when measured from the lateral aspect of the stem.
Table 4.6: Optimal implant angles to reduce calcar stress shielding
Implant Optimal Stem Angle Resulting Head Angle
SMF-Neutral 0.0° 0.0°
SMF-Anteverted 0.0° 6.0°
SMF-Retroverted 15.0° 9.0°
Synergy 10.3° 10.3°
55
To determine the magnitude of this optimization effect, the five levels of anteversion for each implant
were compared using the Tukey test for pairwise comparisons. The interval plot of all four implant cases
is given in Figure 4.8.
Anteversion
Synergy
SMF-Retroverted
SMF-Anteverted
SMF-Neutral
15.00°10.25°
7.50°
3.75°0.00°
15.00°10.2
5°7.50°
3.75°0.00°
15.00°10.25°
7.50°3.7
5°0.00°
15.00°10.25°
7.50°
3.75°0.00°
-20
-30
-40
-50
-60
-70
Av
g %Δ
Me
an
Ca
lca
r S
tre
ss
Figure 4.8: Interval plot (95% C.I.) of percent change in mean calcar stress by implant anteversion
Tukey pairwise comparison showed no significant differences between any anteversion level among
SMF-Neutral, SMF-Anteverted, and SMF-Retroverted implants, except for SMF-Neutral at 15.00°
compared to 0.00° through 7.50° (p < 0.05). Pairwise comparisons could not be conducted for individual
anteversion levels for the Synergy implant because only one experiment was conducted at each level
(compared to five experiments, corresponding to five anterior-posterior positions, for each SMF implant
case; see end of Section 3.3.2). However, a simple comparison of mean stress results in “low”
anteversion (0.00°, 3.75°, 7.50°) against “high” anteversion (7.50°, 10.25°, 15.00°) yielded no significant
difference (p = 0.171).
56
5. Discussion
5.1. Summary of Main Findings
Results from the correlation plots between the FE model and experimental data show good or excellent
correlation for the three broad implant cases (No Implant, SMF, and Synergy). By using slope correction
factors, the FE model can successfully predict experimental results, and can be used for further analysis.
Results from the DOE analysis indicated that calcar stress shielding is present in all implant cases, though
the reduction in stress magnitude for the Synergy implant is approximately double that of the SMF
implants. The peak stress location is shifted distally for all implants, and closely matches each implant’s
stem tip. All peak stress values are below ultimate tensile failure criteria when the femur is in neutral
stance, while they are closer to failure in flexion and above failure in extension. These results indicate
that short stem implants reduce the negative effects of stress shielding compared to long stem implants.
Results from the optimization routine give optimal angles of anteversion to minimize adverse stress
shielding effects for each implant. However, the magnitude of the effect anteversion has on stress
shielding is not significant. This indicates that implant selection is much more important clinically than
anteversion, in reducing the long-term effects of stress shielding.
5.2. Phases 1, 2 Discussion - Correlation of Model to Experiments
Plotting FE model-predicted strains against corresponding experimentally-determined strains showed
strong correlation, as shown in Table 4.1 in the row labelled “Pearson R”: 0.943 (no implant), 0.948
(SMF), and 0.859 (Synergy). These results indicate that the FE model is a strong predictor of
experimental strain, showing a strong linear response.
57
The slopes of the three plots were less than the ideal 1.000 value, which would indicate a 1:1
correspondence between model and experiment. This was accounted for by the slope correction factor,
calculated as the reciprocal of the slope. By multiplying model data by this correction factor, the model
should correspond 1:1 with experimental data. Intercepts were not considered in the data adjustment
since they were over an order of magnitude smaller than the actual strains measured.
This combination of strong correlation and slope correction factors allows the FE model to be sufficiently
adjusted to match experimental data. The model can therefore be used to more precisely and
thoroughly measure the effects of implant orientation and stem length.
5.3. Phase 3 Discussion - Analysis and Optimization
The following sections give discussions of this study’s primary findings, and provide comparisons of each
result to prior studies.
5.3.1. Mean and standard deviation of stress
The results given in Section 4.3.1 are visually illustrated below (Figure 5.1, Figure 5.2). Specifically, the
figures show the distribution of Von Mises stress in the calcar region, coloured by stress magnitude. The
mean stress in the calcar region is significantly reduced in SMF-implanted femurs, and further reduced in
the Synergy-implanted femur (Table 4.2, Figure 4.4). Mean stress is calculated as the average of all nodal
stress values across the region. Mean stress is an indicator of the total load transferred to the region,
normalized to the size of the region. This reduced stress can be seen in the below images as a change in
“average” colour in the calcar region. Reduced stress in a region of bone is the definition of stress
shielding, and results in long-term reduction in bone strength.
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Figure 5.1: Anterior view of calcar Von Mises stress (MPa) in (left – right) Intact, SMF-Neutral, SMF-Anteverted, SMF-Retroverted, and Synergy implanted femurs
Figure 5.2: Medial view of calcar Von Mises stress (MPa) in (left – right) Intact, SMF-Neutral, SMF-Anteverted, SMF-Retroverted, and Synergy implanted femurs
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Similarly, results in Figure 4.5 show the reduction in standard deviation of stress when femurs are
implanted. Standard deviation of stress is also calculated from the nodal stress values in the calcar
region. Standard deviation of stress indicates the “spread” of stress values in the region. This can be
seen in the above images by the reduced range of colours in the calcar region. The “spread” of stress
values indicates the presence or absence of stress concentrations, especially when combined with mean
stress results. Such stress concentrations are important to maintaining bone strength in intact femurs,
since they increase bone strength in critical regions such as the calcar; removing these stress
concentrations leads to weakened bone in critical regions.
Though these mean and standard deviation of nodal stress have not been used in previous literature,
the results they demonstrate have both direct and indirect evidence to support them, particularly mean
stress. Direct evidence in the form of reduced stress in the calcar region has been demonstrated in
previous studies by using experimental testing [61] or FE analysis [63] [64] [65] [77]. Exact values were
rarely given in these publications, so comparisons to the data in Table 4.2 are few.
One study which compared short and long stems [61] showed significant stress shielding only in long
stem implanted femurs, while short stem implanted femurs displayed little change from the intact
femur. No calcar-specific results were explicitly given; additionally, the data was collected from a single
strain gauge in each region, rather than an area within an FE model. Another study in short stem
implants [77] found a 28% change in calcar strain energy density. While strain energy density is not as
commonly used as equivalent strain, these results compare favourably with this study’s results of a
38.97 – 29.87% reduction in calcar mean stress in short stems.
Indirect evidence for stress shielding has been shown by reduced bone mineral density (BMD) in long-
term patients [20] [52]; however, long-term results for short stem implants are not yet available due to
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the concept’s recent introduction. Results from a prospective, randomized trial of four long stem
implants [20] showed peak BMD reductions of 8% – 24% at two post-operative years.
Another trial of a single long stem implant [52] showed 27% BMD reduction in the calcar region, one
year postoperatively. This study also showed that BMD loss is correlated with equivalent stress, as
confirmed in other studies [78] [79]. Therefore, the results from this study may be extrapolated to
predict long term results. The mean calcar stress values in Table 4.2 show that long stem implants
produce significantly higher stress shielding than short stem implants, by factors of 1.6 (Synergy
compared to SMF-Neutral), 2.0 (SMF-Anteverted), and 2.1 (SMF-Retroverted). While the exact
percentage values of change in stress may not produce a 1:1 correlation with changes in BMD, it is
expected that the general result of “short stem implants reduce stress shielding compared to long
implants” will correlate to similar reductions in BMD loss in short stem implants.
It is notable that the SMF-Anteverted and SMF-Retroverted implants created significantly lesser
reductions in calcar stress than the SMF-Neutral implant. All three implants used the same implant
stem, so the only geometric difference was between modular necks. The anteverted and retroverted
necks produce an anterior or posterior head offset of 4mm; additionally, there is a 6mm medial offset in
these two necks, compared to the SMF-Neutral neck. When viewed from above, the SMF-Neutral head
position is 6.8mm from the medial calcar (36.0mm from stem tip, or lateral aspect of the stem); the
SMF-Anteverted and SMF-Retroverted head positions are 13.4mm from the medial calcar (42.2mm from
stem tip/lateral aspect). This increased distance creates a longer moment arm at which the body weight
is applied to the implant head; this is likely the cause of the increased stress, and hence decreased stress
shielding, in the SMF-Anteverted and SMF-Retroverted implants. These conclusions match a previous
study in modular neck selection [80], which found higher strains in longer or anteverted necks than in
shorter or neutral necks.
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5.3.2. Distal load transfer
The third row of results from Table 4.2 gives the distal change in medial peak stress location, measured
axially. These results are visually illustrated in Figure 5.3, along with cutaway images of the SMF and
Synergy stems.
Figure 5.3: Medial view of peak Von Mises stress (MPa) in (left – right): Intact femur, SMF-implanted femurs (SMF-N, SMF-A, SMF-R), and Synergy-implanted femur
The “Max” labels indicate the location of the peak stress, and correspond clearly to the respective data
plot (Figure 4.6). The peak stress location in the intact femur matches results from previous studies [61]
[66] [69]. Peak stress locations in the implanted femurs also match the location of each implant’s stem
tip in the femoral canal; that this match is so clear indicates the significant difference in material
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properties between implant and bone [17]. The stronger stem material absorbs the load and transfers it
distally, rather than evenly throughout the stem length; an implant of ideal material properties would
reduce this distal load transfer [65]. This clear distal load transfer is congruent with previous studies,
which have found distal-dominant load transfer as a result of implant insertion [52] [65]. A more
proximal peak stress location results in greater stress in the calcar region (Figure 5.2), thereby reducing
stress shielding. This is especially true in the distal calcar area, as it is closer to the resulting stress
concentration; however, the proximal region of the calcar remains somewhat stress-shielded. Despite
this, the overall effect is that the short-stemmed SMF implants better reduce the effects of stress
shielding.
It is important to note that the peak stress location indicated for the intact femur in Figure 5.3 (leftmost
image) is not the location of the “true” peak stress. It can be more clearly seen in Figure 5.2 that there is
a location with greater peak stress, 34 mm superior to the indicated peak location in the calcar. This
“true” peak stress location is located more prominently in the femoral neck, a region where stress
shielding is less of an issue for implanted femurs since it is osteotomized. The peak stress location used
for comparisons is located more centrally in the calcar region. However, comparison to the “true” peak
stress location in the femoral neck would yield similar distal shifts, albeit at values 34 mm greater than
in Table 4.2.
5.3.3. Overall peak stress
The peak stress was measured throughout the entire femur. Rather than calculate a percent reduction
from baseline, the actual stress values were recorded, to allow comparisons with failure criteria.
Additionally, stress in each stance was recorded separately, rather than averaged together (Figure 4.7).
The ultimate tensile strength (UTS) of the Sawbones cortical bone is given as 106 MPa [72]. As seen in
the plot, peak stress in neutral stance is below the UTS, while peak stress in flexion is above the UTS in
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several implants, and peak stress in extension is above the UTS in all implants, including the intact
femur. However, these stress levels are a result of the clinical-level load of 3000N, or 4x body weight of
a 75kg patient. This load level is a result of impact loading such as landing from a jump, and is unlikely to
be encountered in extreme flexion or extension stances, let alone in typical patients with hip implants.
Strength values for actual cortical bone compare favourably to the Sawbones synthetic bone. The UTS of
the synthetic cortical bone is 106 MPa [72], while that of actual cortical bone has been measured at 133
MPa [81] through 141 MPa [82]. Compressive strength is generally higher, measured around 200 MPa
[82] or 205 MPa [81]. These values are all given as longitudinal strengths; transverse strength
measurements tend to be lower.
5.3.4. Significance of Implant Orientation and Optimization
The secondary goal of this study was to determine the effect of implant orientation on stress shielding in
the calcar region. The Minitab optimization process yielded one angle of anteversion for each implant
(Table 4.6) which best minimizes changes in stress distribution from the intact femur and reduces peak
stresses to safe levels, across all stances (Section 3.3.4). However, the results of this procedure are not
statistically relevant: pairwise comparison revealed that there are very few statistically significant
differences in stress shielding caused by anteversion, within an implant type. As well, the effect size is
not as large as that of implant selection (Figure 4.8): there is a large degree of overlap in data groups of
the same implant. Levels of anteversion for the Synergy stem could not be compared as-is, since there
was a single experiment performed at each anteversion level (see end of Section 3.3.2). However, when
grouped into low and high-anteversion data groups, there was no significant difference between the
groups. These conclusions match a previous study on implant orientation in short stem implants [77],
which concluded that implant orientation had no significant effect on cortical strains.
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5.4. Clinical Implications and Future Directions
The results of this study have immediate clinical implications. While stress shielding has been previously
investigated, this study determined the specific amount of stress shielding caused by several implants.
These results can help clinicians make informed decisions regarding implant selection, both regarding
these specific implants, and short and long stem implants in general. Specifically, these results
demonstrate that the short stem SMF implant – with three of its various modular necks – produces
stress shielding at about half the magnitude of the long stem Synergy implant. Modular necks that
increase the implant head’s distance from the stem – in this study, the SMF-Anteverted and SMF-
Retroverted necks – further reduce stress shielding, though by a smaller margin. These results are
particularly important for younger patients, who will benefit more from long-term analyses of bone
integrity. Additionally, these predictions of long-term stress shielding are helpful because there are
currently no long-term analyses of short stem implants. The stress results here could also be correlated
with BMD measurements, as previously investigated [52] [78] [79]; if successful, this correlation could
yield more accurate predictors of long-term bone integrity.
Despite the improved long-term prospects of short stems and long necks, it must be cautioned that too
high a stress level, especially in the calcar, may result in short-term destabilization and micromotion of
the implant. Short term analyses, especially comparing various neck lengths, should be performed to
determine whether the increased stress in the calcar is purely beneficial, or may increase the risk of
micromotion. Such short-term analyses are detailed in Section 5.5.2.
The distal shift demonstrated also has possible clinical implications. By quantifying how the calcar stress
concentration is shifted distally, it is evident that this stress location is linked to the implant geometry.
This could allow clinicians to select an implant that distributes stress according to the patient’s needs:
for example, patients with pre-existing bone weakness, perhaps due to prior injury, at the location of
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the SMF stress concentration (Figure 5.3) may not be appropriate for this implant, to prevent additional
stress to this area. In contrast, additional stress to this area might be beneficial in helping the healing
process; further investigation may be warranted to determine whether this implant-induced stress
concentration would be beneficial or harmful.
The peak stress results (Figure 4.7) do not provide any new clinical implications of importance. In neutral
stance, peak stress is reduced for implanted femurs compared to the intact femur, so cortical fracture is
unlikely in this stance once the implant is fully stable and osseointegrated. In flexion, peak stresses
increased somewhat in implanted femurs, with some stresses above the UTS of 106 MPa. This would
contraindicate its use with high impact activities in this stance such as jumping or skiing. This is
congruent with post-operative guidelines which recommend low impact activities [83]. Peak stress in
extension showed the largest increase over intact femurs. However, as discussed in Section 5.3.3, high
impact loads are unlikely to be encountered at the angle of extension tested. Some exercises in strength
training or yoga may encounter such angles; however, such activities are typically advised against
because of the possibility of dislocating the joint [84].
The results presented can also alleviate some concerns about optimal implant orientation: the degree of
anteversion has no significant effect on stress shielding, regardless of the implant selected. However,
these results are only applicable to a specific range of anteversion: a 0°-15° anteversion range (in
addition to the 12° angle of torsion, or the femur’s natural anteversion with respect to the knee) was
pre-selected, and the maximum anterior-posterior positions were visually estimated to keep the implant
within the cortical bone. FE simulations were not run outside this “safe” range. As discussed, there is a
large margin of error for potential surgeon over and under-estimation of implant anteversion (Section
2.5), so care must be taken to keep the implant within this safe range. The present FE model could be
adapted to include anteversion levels outside this safe range; it is likely that these simulations would
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result in cortical failure since the implant would significantly impinge on the cortical bone. Furthermore,
the varus-valgus angulation of the implant was not investigated, but was fixed as indicated by the SMF
surgical technique [22]. Short stems have a larger margin of varus-valgus implantation error because the
shorter length allows greater freedom to rotate within the femoral canal. Additionally, the uncemented
nature of this stem means that positioning errors are more difficult to fix intraoperatively. The FE model
could be adapted to analyze varus-valgus malposition, to better understand the associated short and
long-term effects.
The FE model can be adapted to analyze any other implant under similar conditions. The large number
of orientation parameters, plus the DOE approach used, could easily determine the stress response to a
wide variety of changes in implant positioning. Another parameter that might be considered is bone
quality: while the current study provides results pertaining to a healthy bone stock, an osteoporotic
patient may have different implant design or positioning requirements. The DOE approach could be
used to improve existing implant designs, or to develop entirely new implants. An FE model has been
used to develop an implant stem customized to a specific population [65]; using the DOE method from
the present study would allow a more comprehensive analysis of implants for other patient populations.
It is important to caution that these results are solely indicative of long-term stress shielding in the
bone. The analysis assumed fully bonded contact between the implant and bone. This is not the case
during the period immediately following implantation: the bone has not fully osseointegrated into the
porous implant coating, meaning that the implant is relatively free to shift. Prior studies have
investigated implant migration using radiostereometric analysis (RSA) [85] and FE analysis using
frictional implant-bone contact [86] [87]; the current FE model could be adapted to model short-term
migration using the conditions described in the indicated studies. However, short-term results for short
stem implants in general have shown good results (Section 1.1.3), so such an adaptation to the model
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would only serve to quantify this migration. Modelling short-term stability using FE models will be
further discussed in Section 5.5.2.
One possible disadvantage to greatly reducing stress shielding is the loss of a clinical indicator of
osseointegration. When an implant is fully osseointegrated and stable, some stress shielding will occur
as stress is transmitted through the implant-bone interface. If an implant is not osseointegrated, no
stress shielding will occur, though this may ultimately lead to more significant aseptic loosening.
Therefore, some degree of stress shielding is beneficial as an indicator of proper fixation. Stress
shielding is typically measured clinically using x-rays, manifested visually as a change in cortical thickness
in the calcar from preoperative images, or when compared to the non-operated side [88]. Stress
shielding can also be quantified as a change in bone mineral density (BMD) using computed
tomography-assisted osteodensitometry [89] or dual-energy X-ray absorptiometry (DEXA) [52]. The
greatest reduction in calcar BMD is typically seen two years post-operatively [20]. If stress shielding is to
be further reduced through changes in implant design, alternate methods of confirming full fixation
should be investigated that can be applied to any patient. Currently, radiostereometric analysis (RSA)
can be used to measure migration with a high degree of accuracy [85]; however, RSA requires
radiopaque markers to be placed in the patient at the time of surgery, and is thus not applicable for all
patients. An alternative would be to investigate what the optimal amount of stress shielding is, which
both maximizes remaining calcar bone strength, yet remains detectable by clinicians using existing
methods as a sign of fixation.
There are other factors besides stem length that contribute to reducing stress shielding, and could not
be controlled for in this study. Broadly speaking, the cross-sections of the SMF and Synergy stems are
rectangular and circular respectively. Not only does the cross-sectional area affect the implant’s fit and
fill within the canal [65], but also the flexural stiffness and resistance to torsion of the stem [88]. A
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greater medial-lateral width increases implant stiffness in the frontal plane, such as in axial compression
in neutral stance; however, more flexible stems can better match the mechanical properties of bone, so
a greater implant stiffness would result in increased stress shielding. In contrast, greater medial-lateral
width increases stability in resistance to axial torsion, such as in stair-climbing; this stability is important
in the short term as the implant is becoming osseointegrated. To fully appreciate the effect of stem
length on stress shielding, identical implants should be modelled with only the stem length differing. The
amount of taper in the stem – full length, partial length, none – could also be modelled and controlled
for.Another implant design factor affecting stress shielding is the implant’s material strength. By
decreasing the material’s Young’s Modulus, a higher flexibility is achieved, and thus a closer match to
the material properties of bone; however, this increased flexibility may result in greater micromotion as
the implant is permitted more freedom of movement within the canal, ultimately leading to loosening
[17]. Finally, the type of bond-coating and area of application affects uncemented implant design.
Hydroxyapatite coating has been shown to increase pull-out strength of hip stems [90], but it does not
improve overall stem survival rates compared to non-HA coatings such as porous or sand-blasted stems
[91]. Stems are typically coated proximally to improve osseointegration in the metaphysis, while left
uncoated distally to ease insertion and allow some motion during early osseointegration; shorter stems
may require less non-coated surface to improve early osseointegration. The material stiffness will also
affect the amount of bond-coating applied, and is worth investigating further.
5.5. Limitations and Sources of Error
5.5.1. Phase 1 – Mechanical Testing
This study has a number of limitations which limit the potential clinical implications of the results. The
simplified test setup does not account for a number of muscle attachments which would create
additional forces in the femur; among these muscles are the vastus medialis, vastus lateralis, tensor
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fascia lata, ilio-tibial tract, and piriformis. These forces have been replicated in studies that use FE
models alone [64] [65]; however, the experimental nature (Phase 1) of this study makes this more
difficult to accomplish. One study which combined mechanical and FE testing [63] developed a load jig
which uses a pulley and lever system to simulate the hip abductors. Adding a similar setup to the current
study is feasible but would require a more realistic implant procedure: a bandsaw was used to cut
straight through the entire greater trochanter along the osteotomy plane, because these muscle
attachments were not planned. This would likely produce a more physiological loading condition, and
deserves future investigation.
While using a single, consistent femoral geometry has advantages such as reproducibility and
predictability, it limits the applicability of the results to femurs with similar geometry. Studies have
demonstrated how implant stems are designed for certain populations, typically European and North
American [92] [93]. Patient-specific implant stems have been investigated and shown promising results,
especially in improving fit and fill within the femoral canal for non-European or North American
morphologies [65]. While results customized to individual patients or morphologies are ideal, the results
from the current study indicate overall trends in stress shielding as a result of substantial changes in
stem geometry.
Another disadvantage to using synthetic femurs is their overall homogenous, isotropic geometry. This is
especially true for the Sawbones synthetic cancellous bone, which is a solid foam matrix. In contrast, the
cancellous matrix in human bones is heterogeneous and anisotropic: as previously stated (Section 2.2.1),
three series of trabeculae (arcuate, medial, and lateral systems) are oriented to withstand compressive
forces. Such anisotropic anatomy may be cumbersome to produce, and likely would not yield
significantly different results since cortical bone strength is an order of magnitude higher.
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The femurs were potted in cement before implantation. The orthopedic resident noted that the
cancellous bone in the synthetic femur was stiffer than that of a typical patient. Since the SMF implant
was uncemented, a large amount of hammering was required to broach the cancellous bone and seat
the implant. All six synthetic femurs implanted with the SMF stem separated from the concrete base
(the canal for the Synergy stem was opened primarily through drilling, with minimal broaching). The
concrete retained its shape, preventing upward displacement and rotation; however, downward
displacement of the femur was possible, so a steel plate was inserted below the femur to reduce this
displacement. This possibility of initial displacement may have produced inaccurate strain results, but
the good correlation results make this unlikely. For future studies, it is recommended to insert the
implants before potting the femurs, if possible.
It may have been advantageous to place more strain gauges in the calcar region. The number of strain
gauges was limited by the data acquisition hardware to eight channels; however, one rosette gauge
required three channels and could be replace by three single-channel linear gauges. By using more strain
gauges in the calcar region, a more accurate representation of stress shielding in the calcar could be
obtained. The strain gauges used in this study were spread out across the femur’s length to obtain a
broad representation of stress in the entire femur, such as quantifying the distal shift in medial stress
concentration. Now that this has been measured, future studies could focus on a more precise
quantification of stress shielding specific to the calcar region.
A single slope correction factor (Table 4.1) was calculated based on data from all six strain gauges in all
test conditions. As seen in Figure 4.1 through Figure 4.3, the data did not fall neatly into a single line-of-
best-fit. It may be that different strain gauges had different correlation slopes, and would therefore
require different slope correlation factors. This may improve accuracy in these discrete points on the FE
model; however, it would be difficult to apply multiple correction factors to entire regions of the femur
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(e.g. slope correction factor of 1.3 in the calcar, 1.1, in distal shaft, etc) without defining strict
boundaries. One potential solution would be to weight the single slope correction factor proportionately
to regions of interest (i.e. calcar strain is more important than distal shaft strain, so the entire model’s
slope correction factor would be closer to the calcar’s factor than the average of all correction factors).
This would provide more accurate quantification of stress in the calcar, while maintaining a single,
simple model of stress in the entire femur.
While the unimplanted and Synergy FE models each had their own corresponding slope correction
factor, all three SMF implant types – Neutral, Anteverted, and Retroverted, shared a single correction
factor. This was done out of convenience, since the factors for each individual implant type were similar
in magnitude. While this is acceptable when comparing the overall stress shielding effect between
unimplanted, SMF-implanted, and Synergy-implanted femurs, comparisons between these three SMF
modular necks is less accurate. Separate correction factors for each neck type would improve this
accuracy.
Inaccuracies in strain gauge placement were inevitable, but reduced as much as possible. They were
placed by hand using ruler measurements. The FE model included parameters to adjust the virtual strain
gauges’ positions, which were compared to photographs to obtain accuracy to within 1 mm. While this
level of accuracy is fairly high, it may have resulted in weaker correlation between mechanical and FE
results. A measurement jig may help place strain gauges more accurately, but it is possible that there
would be no significant benefit – the correlation coefficients were strong regardless.
The actual test method represents only a fraction of scenarios encountered by patients. Other studies
have tested torsional strength of synthetic and cadaveric femurs [54]. High levels of torsion are unlikely
to be encountered in hip replacement patients as high-intensity activities are discouraged; additionally,
a number of different, potentially complex loading jigs would be required to test torsion for both intact
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and implanted femurs. Similarly, high levels of flexion, abduction, and adduction could be tested to
reveal the response to cantilevered bending; however, such extreme hip angles are unlikely in patients
due to the likelihood of dislocation [84]. However, further testing could reveal how the implanted femur
responds to atypical situations, such as falls. Such testing could also utilize destructive testing – avoided
in this study because the specimens were reused many times, and only typical “everyday” situations
were considered.
In order to test such atypical conditions, the physical constraints of the test setup must be considered.
Primarily, the Instron 8874 testing machine used has a base fixation plate which only allowed femur
angles up to the 15° tested in this study. Higher angles, such as 90° of flexion seen in a squatting
position, would require a modified fixation jig to mount the femurs securely, or a shorter working length
of femur. While fatigue testing through cyclic loading is possible with this system, an alternate method
would be necessary to investigate impact loads.
5.5.2. Phases 2, 3 – Finite Element Model
Another limitation of this study is the exclusive use of bonded contacts in the FE model. Bonded contact
was chosen for this study after initial attempts with frictional contact resulted in excessively long
solution times and frequent software crashes; additionally, the study method required a high number of
experiments, which emphasized a fast solution time. Previous studies have used bonded contacts for
similar purposes [17] [65], while other studies have used frictional contacts instead [63] [64] [87].
Frictional contact has been used to study micromotion, an indicator of short-term stability [86] which
could complement this study’s long-term stress shielding results. Another study has used a combination
of bonded and low-friction contacts [77]: the bonded contact simulated full osseointegration in the
proximal porous stem coating, while the low-friction contact was used for the smooth stem tip. A
smaller area for frictional contact reduces the solution time. However, full osseointegration assumes
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that the implant survived the initial short-term period of developing stability; therefore, bonded contact
would not be helpful in determining the short-term effects. Frictional contacts could be employed in
future studies, as long as the number of required solutions is reduced.
An alternative to frictional contacts in predicting implant micromotion is to analyze interfacial shear
stresses in the linear, bonded-contact model. Specifically, shear stress in the axial direction is strongest
at the proximal bone-implant interface, and causes micromotion slippage of the implant with respect to
the surrounding bone [94] [95]. Shear stress in the model can be compared to the normal stresses using
an appropriate frictional coefficient to determine whether micromotion would occur. This eliminates the
non-linear approach of frictional models, significantly reducing solution time, while predicting the
immediate micromotion that would be obtained from such models. The drawback is that only initial
micromotion would be measured; using a non-linear frictional model, the implant’s overall micromotion
over time, such as how it “settles” into the bone, could be measured.
One potential source of error that is readily visible is small geometric imperfections in the SMF-
Anteverted and SMF-Retroverted implant FE models. In Figure 4.7 and Figure 4.8, these two implants
show much larger ranges of stress than the other implants. Other result plots show singularities
(asterisks) outside the range of most data points. These singularities are areas of abnormally high stress
concentrations from the FE simulations. They are caused by miniscule flaws in the implant FE geometry,
which were eliminated as much as possible but not entirely; these flaws were more prevalent in the
SMF-Anteverted and SMF-Retroverted implants than the SMF-Neutral and Synergy implants. Typically,
one in four DOE simulations would result in peak stresses several orders of magnitude higher than
expected, usually at a single node. The DOE input parameters would be tweaked slightly and the
simulation rerun, until the peak stress fell under a normal order of magnitude. Despite this iterative
tweaking, there were still some abnormal stress values that could not be eliminated.
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It is possible that some locations of stress concentration were not abnormal. These areas may indicate
locations of microfractures in the cortical bone, which could help predict sites where cracks may initiate.
These locations may also show where the implant design could be improved to reduce stress
concentrations. These stress concentrations could be addressed by refining the mesh in the surrounding
area, or by improving the initial FE implant model. Fatigue analysis may provide further insight as to
whether these areas of stress concentration, whether in the bone or implant, would be dangerous.
Another potential source of inaccuracy between the experimental specimens and FE simulation is the
number of parameters used in the FE model. The seven parameters used are given in Section 3.2.2. Two
parameters (working length and adduction) affect only the femur; two parameters (osteotomy angle
and osteotomy level) affect both the femur and implant position; three parameters (anteversion,
anterior-posterior position, and medial-lateral position) affect only the implant position. The implant’s
superior-inferior position is linked to the osteotomy level, while the implant’s varus-valgus angle is
linked to the osteotomy angle. While it was assumed that the implant was fully seated in the resected
bone, a higher degree of accuracy may be possible by creating separate parameters for the femur and
implant; additionally, the effect of improper varus-valgus angle or inadequate implant seating depth
could be investigated using these extra parameters. Two other parameters could be possible: femoral
anteversion about its shaft (rotation in the transverse plane), and implant flexion and extension
(rotation in the sagittal plane). Minor variations were observed in these two degrees of freedom in the
experimental specimens; however, they were not judged significant enough to warrant additional
measurements in the FE model. These two parameters would be held constant in the final “Analysis”
phase, similar to the five other parameters held constant for the Design of Experiments setup; this
further reduces the need for these two parameters.
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6. Conclusion
This study has demonstrated how implant stem length and orientation produce stress shielding in the
calcar. Two implant stems, short and long, were compared to an intact femur to determine the relative
change in equivalent stress in cortical bone. Three modular necks were used in the short stem implant,
further investigating the effect of implant neck design. Using a combination of experimental testing and
finite element simulation, the extent of stress shielding in the calcar was measured and analyzed for all
physiologically possible implant positions. Results showed stress shielding and distal load transfer
present for all implants; the long stem implant displayed significantly higher stress shielding and distal
load transfer than the short stem implant; the anteverted and retroverted modular necks performed
slightly better than the neutral neck. Implant orientation had no significant effect on stress shielding.
This study has both investigated the effect of implant selection and orientation on stress shielding, and
developed a viable model which can be easily adapted to investigate a variety of other factors related to
hip implants.
76
7. References
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8. Appendix
The following code was written for use as an Ansys “command” in the Mechanical environment. Three
named selections are created by selecting geometry as shown in Figure 3.9 and naming the selections as
follows: “calcar_selection” (calcar), “max_selection” (medial femur), and “strain_selection” (overall
osteotomized femur). One input, “ARG1” is used as a stress multiplier, as given by the slope correction
factors (Table 4.1). A number of outputs are given: OVR_mn, OVR_sd, CAL_mn, and CAL_sd calculate
mean and standard deviation of nodal stress in the overall femur and calcar respectively;
OVR_MaxStressVal and MEDMaxStressVal calculate peak stress in the overall and medial femur;
MEDMaxStressLocSI and MEDMaxStressLocAP calculate the superior-inferior and anterior-posterior
location of the peak stress.
! initialization ##########################################################################
resume,,db ! reopens MAPDL solution .DB file
set,last ! use data from final timepoint
! calculate overall mean ##################################################################
cmsel,s,strain_selection,elem ! select "strain_selection"
*get,ndnum,node,0,count ! count all nodes -> "enum"
! initialize variables
NN = 0 ! element counter
OVR_mn = 0 ! stack for overall mean vonmises stress
prev_str = 0
valid_nodes = 0
*do,i,1,ndnum ! loop through all nodes
NN = ndnext(NN) ! get number of next element
*get,nd_str,node,NN,s,eqv ! vonmises stress of node NN -> "nd_str"
*if,nd_str,ne,prev_str,then
OVR_mn = OVR_mn + nd_str ! add current node's stress to stack
valid_nodes = valid_nodes + 1
*endif
prev_str = nd_str
*enddo
OVR_mn = OVR_mn/valid_nodes ! mean stress
! calculate overall std dev ###############################################################
NN = 0 ! re-initialize element counter
OVR_sd = 0 ! stack for SD vonmises stress
86
meandiff = 0 ! temp variable for (Xi - Xmean)
*do,i,1,valid_nodes ! restart loop through all elements
NN = ndnext(NN)
*get,nd_str,node,NN,s,eqv ! vonmises stress of node NN -> "nd_str"
meandiff = nd_str - OVR_mn ! calculate (Xi-Xmean)^2
meandiff = meandiff*meandiff
OVR_sd = OVR_sd + meandiff ! add to stack
*enddo
OVR_sd = sqrt(OVR_sd/valid_nodes) ! calculate SD
my_OVR_mn = OVR_mn*ARG1 ! multiply mean by correction factor
my_OVR_sd = OVR_sd*ARG1 ! multiply sd by correction fact
! #########################################################################################
! calculate calcar mean ###################################################################
cmsel,s,calcar_selection,node ! select "calcar_selection"
*get,ndnum,node,0,count ! count all nodes -> "ndnum"
! initialize variables
NN = 0 ! node counter
CAL_mn = 0 ! stack for overall mean vonmises stress
prev_str = 0
valid_nodes = 0
*do,i,1,ndnum ! loop through all nodes
NN = ndnext(NN) ! get number of next element
*get,nd_str,node,NN,s,eqv ! vonmises stress of node NN -> "nd_str"
*if,nd_str,ne,prev_str,then
CAL_mn = CAL_mn + nd_str ! add current node's stress to stack
valid_nodes = valid_nodes + 1
*endif
prev_str = nd_str
*enddo
CAL_mn = CAL_mn/valid_nodes
! calculate overall std dev ###############################################################
NN = 0 ! re-initialize element counter
CAL_sd = 0 ! stack for SD vonmises stress
meandiff = 0 ! temp variable for (Xi - Xmean)
*do,i,1,valid_nodes ! restart loop through all elements
NN = ndnext(NN)
*get,nd_str,node,NN,s,eqv ! vonmises stress of node NN -> "nd_str"
meandiff = nd_str - CAL_mn ! calculate (Xi-Xmean)^2
meandiff = meandiff*meandiff
CAL_sd = CAL_sd + meandiff ! add to stack
*enddo
87
CAL_sd = sqrt(CAL_sd/valid_nodes) ! calculate SD
my_CAL_mn = CAL_mn*ARG1 ! multiply mean by correction factor
my_CAL_sd = CAL_sd*ARG1 ! multiply sd by correction fact
! #########################################################################################
! calculate overall max stress value ######################################################
cmsel,s,strain_selection,node ! select nodes in "max_selection"
*get,ndnum,node,0,count ! count all nodes -> "ndnum"
! initialize variables
NN = 0 ! nodal counter
my_OVRMaxStressVal = 0 ! value of max stress
tempStress = 0 ! value of current node's stress
tempLoc = 0 ! value of current max-stress node's SI-loc
*do,i,1,ndnum
NN = ndnext(NN)
*get,tempStress,node,NN,s,eqv ! get current node's stress
*if,tempStress,GT,my_OVRMaxStressVal,then
my_OVRMaxStressVal = tempStress
*endif
*enddo
my_OVRMaxStressVal = my_OVRMaxStressVal*ARG1
! #########################################################################################
! calculate medial max stress location ####################################################
cmsel,s,max_selection,node ! select nodes in "max_selection"
*get,ndnum,node,0,count ! count all nodes -> "ndnum"
! initialize variables
NN = 0 ! nodal counter
my_MEDMaxStressVal = 0 ! value of max stress
my_MEDMaxStressLocSI = 0 ! SI location of max stress
my_MEDMaxStressLocAP = 0 ! AP location of max stress
tempStress = 0 ! value of current node's stress
tempLoc = 0 ! value of current max-stress node's SI-loc
*do,i,1,ndnum
NN = ndnext(NN)
*get,tempStress,node,NN,s,eqv ! get current node's stress
*if,tempStress,GT,my_MEDMaxStressVal,then
*get,tempLoc,node,NN,loc,y
! *if,tempLoc,LT,ARG2,then
my_MEDMaxStressVal = tempStress
my_MEDMaxStressLocSI = tempLoc
*get,my_MEDMaxStressLocAP,node,NN,loc,z
! *endif
*endif
*enddo
my_MEDMaxStressVal = my_MEDMaxStressVal*ARG1
! END OF CODE ##########################################################################