ArticlePDF Available

Joint Load Considerations In Total Knee Replacement


Abstract and Figures

Estimates of knee joint loadings were calculated for 12 normal subjects from kinematic and kinetic measures obtained during both level and downhill walking. The maximum tibiofemoral compressive force reached an average load of 3.9 times body-weight (BW) for level walking and 8 times BW for downhill walking, in each instance during the early stance phase. Muscle forces contributed 80% of the maximum bone-on-bone force during downhill walking and 70% during level walking whereas the ground reaction forces contributed only 20% and 30% respectively. Most total knee designs provide a tibiofemoral contact area of 100 to 300 mm2. The yield point of these polyethylene inlays will therefore be exceeded with each step during downhill walking. Future evaluation of total knee designs should be based on a tibiofemoral joint load of 3.5 times BW at 20 degrees knee flexion, 8 times BW at 40 degrees and 6 times BW at 60 degrees.
Content may be subject to copyright.
VOL. 79-B, N
. 1, JANUARY 1997 109
M. S. Kuster, MD, Orthopaedic Surgeon
A. G¨achter, Professor and Chairman
Klinik f¨ur Orthop¨adische Chirurgie, Kantonsspital, 9007 St Gallen,
G. A. Wood, PhD, Associate Professor
Neuromuscular Performance Laboratory, Department of Human
G. W. Stachowiak, PhD, Associated Professor
Tribology Laboratory, Department of Mechanical and Materials
The University of Western Australia, Nedlands, Australia 6907,
Correspondence should be sent to Dr M. S. Kuster.
©1997 British Editorial Society of Bone and Joint Surgery
0301-620X/97/16978 $2.00
From the University of Western Australia, Australia
Estimates of knee joint loadings were calculated for 12
normal subjects from kinematic and kinetic measures
obtained during both level and downhill walking. The
maximum tibiofemoral compressive force reached an
average load of 3.9 times body-weight (BW) for level
walking and 8 times BW for downhill walking, in each
instance during the early stance phase. Muscle forces
contributed 80% of the maximum bone-on-bone force
during downhill walking and 70% during level walking
whereas the ground reaction forces contributed only
20% and 30% respectively.
Most total knee designs provide a tibiofemoral
contact area of 100 to 300 mm
. The yield point of these
polyethylene inlays will therefore be exceeded with each
step during downhill walking. Future evaluation of total
knee designs should be based on a tibiofemoral joint
load of 3.5 times BW at 20° knee flexion, 8 times BW at
40° and 6 times BW at 60°.
J Bone Joint Surg [Br] 1997;79-B:109-13.
Received 6 June 1996; Accepted 24 September 1996
Several recent studies have reported severe wear of poly-
ethylene tibial components.
The long-term problems
associated with joint wear debris, such as loosening and
infection, are well known. Wear is dependent on a number
of factors including contact area, load, material properties,
thickness of the polyethylene inlay and the length of time
that the component has been implanted.
The most
destructive wear process is fatigue, which occurs through
repeated high loads and cyclic stressing. Load is dependent
both on physical activity and on body-weight.
The increasing long-term successes being achieved with
total knee replacement means that younger, and consequently
more active, patients are being treated. This places an
increased mechanical demand on the prosthesis which exceeds
the design limits of many of the currently used devices.
The moments and forces about the knee vary substan-
tially for different daily activities.
Biomechanical studies of
knee joint loading have consistently estimated maximum
joint compressive forces to be about 4 to 4.5 times body-
weight during daily activities.
This range of values has
become a design criterion for most currently used knee
prostheses, but recent studies have indicated that loadings
can be much higher even during level walking.
finding is consistent with the increasing incidence of reports
of severe wear in joint replacements.
We present
quantitative joint load data and suggest new criteria for use
in the biomechanical evaluation of total knee prostheses.
We obtained estimates of knee joint loading for 12 normal
subjects (6 male and 6 female) ranging in age from 23 to 37
years (mean 27.9), in height from 158 to 187 cm (mean
171) and in weight from 49 to 90 kg (mean 70.8). Reflec-
tive markers were located superficial to the 5th metatarso-
phalangeal, ankle, knee and hip joints. Spatial trajectories
were recorded using a video-based motion analysis system
with two cameras sampling at 60 Hz (APAS, Ariel Dynam-
ics, Inc, Trabuco Canyon, California) whilst the subjects
walked across a level floor and down a purpose-built ramp
of 19% gradient. Ground reaction force data were simulta-
neously obtained from a Kistler force platform (Type
9281B, Winterhur, Switzerland). Ground reaction forces
during downhill walking were measured using an alumin-
ium plate bolted to the force platform. Step frequency for
both downhill and level walking was controlled by means
of a metronome set at 120 steps/min.
The marker trajectories in the sagittal plane were smoo-
thed using a Butterworth 4th-order low-pass digital filter
with a cut off frequency of 7 Hz prior to the derivation of
segmental orientations and centre of mass locations. Time
derivatives of these measures were then calculated by finite
differences. Finally, planar joint reaction forces and net
joint moments at the ankle, knee and hip were estimated
from these kinematic data and force platform measures
using standard inverse dynamics procedures and anthropo-
metric values from Winter.
A knee joint model, previously described by Nisell,
was then used to calculate the tibiofemoral (bone-on-bone)
force from the mean joint reaction forces and knee extensor
moments which had been derived from the inverse dynamic
analysis. The total uncertainty based on the standard error
of the mean of the peak bone-on-force was calculated as
outlined by Campion.
Full details of the instrumentation
used can be found in the authors’ previously published
The highest knee joint loadings occurred during downhill
walking. The peak joint moments occurred at 41 ± 6° knee
flexion and were 2.75 ± 0.5 Nm/kg for females and 2.70 ±
0.7 Nm/kg for males. The vertical joint reaction forces
were 15.2 ± 1.6 N/kg for females and 15.5 ± 1.9 N/kg for
males. The kinetic and kinematic data for males and
females were not significantly different (p > 0.85). The
joint model used in this investigation defined the lever arm
of the quadriceps muscles of female subjects as being
significantly smaller than that of male subjects. Thus the
actual tibiofemoral model predictions were consistently
smaller for male subjects. The peak tibiofemoral force for
male subjects was 7 times BW during downhill walking,
whereas it reached 8 times BW for female subjects. Values
obtained for level walking (at 20° knee flexion) were
approximately 50% of those for downhill walking giving
values of 3.4 times BW for male and 3.9 times BW for
female subjects. The standard error of the mean in the
prediction of peak force was calculated to be 13% for level
and downhill walking. The mean female knee joint com-
pressive forces for the duration of the support phase during
downhill walking and level walking are shown in Figure 1.
The bone-on-bone compressive forces are shown as are the
muscle and gravitational force (ground reaction force) con-
tributions to this load. As can be seen from Figure 1,
extensor muscle force is by far the greatest contributor to
the joint compressive force during level (70%) and down-
hill (80%) walking.
From the estimated knee joint loads, we calculated the
stress on the tibial plateau for a 70 kg female in several
walking tasks, plotted against varying tibiofemoral contact
area, assuming a uniform pressure distribution. The results
are shown in Figure 2, together with an indication of the
yield range for ultra high molecular weight polyethylene
(UHMWPE). It is clear that in order to obtain stress levels
which are safely below the yield point of UHMWPE for all
Fig. 1
Mean tibiofemoral joint loadings of the six female subjects during the stance phase in level and downhill walking. The values are
reported in multiples of body-weight (BW) and normalised to 100% of the stance time. Heel strike occurs at 0% and toe-off at
100%. The total uncertainty (SEOM) is indicated for the peak tibiofemoral compressive force values.
daily activities a contact area greater than 400 mm
This study demonstrates that the muscle forces contribute
80% of the maximum bone-on-bone force during downhill
walking and 70% of the maximum bone-on-bone force
during level walking. The magnitude of the ground reaction
force is not the best predictor of the joint load, the muscle
moments are more reliable. These patellar ligament forces
are based on calculations of the net muscle moment of
force acting about the knee, which is a limitation of the
inverse dynamics approach. Quadriceps effort required to
overcome antagonistic effects of the knee flexor muscles is
not included. Electromyographic activity, recorded from
our subjects during their downhill walking, clearly indicat-
ed the presence of hamstring, gastrocnemius and quad-
riceps muscle co-activity during the stance phase.
measures of joint compression reported here are therefore
conservative estimates but still exceed eight times body-
weight for downhill walking.
Force values equivalent to three to four times BW have
previously been used in most biomechanical tests evaluat-
ing total knee replacements.
Estimates of the tibiofe-
moral bone-on-bone forces in our study were close to four
times BW even during level walking and more than eight
times BW during downhill walking. Collins
calculated the
knee joint loads during level walking using an optimisation
method which incorporated muscle coactivation of agonists
and antagonists. He concluded that the tibiofemoral loads
range from 3.9 to 6.0 times BW. Jefferson et al
found that
the maximum tibiofemoral loads are up to 6.3 times BW,
while Wyss et al
report values ranging from 2.5 to 5 times
BW. Our results for level walking are well within the limits
of these predictions. As the experimental set-up for level
and downhill walking did not change, the calculated loads
for downhill walking, allowing for the angulation of the
plate, present a valid comparison.
Loads for biomechanical evaluation of patellar compo-
nents have been considered to be in the range of 0.7 to 2
times BW.
Some researchers, in order to evaluate different
designs, assumed loads as low as 0.15 times BW for level
walking and 2.2 times BW for walking downstairs.
research indicates higher loadings; for level walking patello-
femoral joint forces of 1.3 to 1.8 times BW,
for down-
stairs walking 5.5 times BW
and for downhill walking 5 to
7 times BW have been suggested.
For some sports activ-
ities such as jumping
or weightlifting
the loads imposed
on the patellofemoral joint are close to 20 times BW.
A review of the available literature indicates that the
majority of authors use the lowest tibiofemoral and patello-
femoral joint loads reported for the evaluation of contact
stresses in total joint replacement.
In order to
improve the design of total knee replacements, it is neces-
sary to adopt higher tibiofemoral and patellofemoral loads.
Contact area has a very profound effect on joint stress
(Fig. 2). The reported average contact area of a natural knee
joint ranges from 765 mm
to 1150 mm
After complete
medial and lateral meniscectomy the tibiofemoral contact
area is approximately 520 mm
, depending on the load.
Assuming a uniform load distribution and a load of eight
times BW the estimated stress on the articular cartilage is
only about 10 MPa for a knee joint without menisci and less
than 5 MPa for a healthy knee. The contact area of most
total knee prostheses is between 80 and 300 mm
on the load, flexion angle and design,
leading to contact
stresses on the UHMWPE inlay as high as 60 MPa; this
exceeds the yield point of 20 MPa for UHMWPE.
A contact area of approximately 400 mm
is necessary to
avoid stresses to the polyethylene inlay that are above the
yield point of 20 MPa. This contact area should be main-
tained throughout a flexion range of 0° to 60° to accom-
modate the high loads of downhill and downstairs walking.
Congruent prostheses significantly reduce polyethylene
wear, and line or point contact should be avoided.
In summary, the design of knee replacements should
allow for the much higher joint loadings now being esti-
mated through gait analysis if severe wear is to be
VOL. 79-B, N
. 1, JANUARY 1997
Fig. 2
Plot of the tibial plateau stress versus the contact area for different daily
activities for a 70 kg female subject. A uniform stress distribution is
assumed. The range of tibiofemoral contact area of current knee pros-
(TK), a knee joint after meniscectomy
(MK) and a natural knee
(NK) are shown. The horizontal bar indicates the yield range of
No benefits in any form have been received or will be received from a
commercial party related directly or indirectly to the subject of this
1. Engh GA, Dwyer KA, Hanes CK. Polyethylene wear of metal-
backed tibial components in total and unicompartmental knee pros-
theses. J Bone Joint Surg [Br] 1992;74-B:9-17.
2. Engh GA. Failure of the polyethylene bearing surface of a total knee
replacement within four years: a case report. J Bone Joint Surg [Am]
3. Lindstrand A, Ryd L, Stenstr¨om A. Polyethylene failure in two total
knees: wear of thin, metal-backed PCA tibial components. Acta
Orthop Scand 1990;61-:575-7.
4. Jones SMG, Pinder IM, Moran CG, Malcolm AJ. Polyethylene
wear in uncemented knee replacements. J Bone Joint Surg [Br]
5. Bartel DL, Bicknell VL, Wright TM. The effect of conformity,
thickness and material on stresses in ultra-high molecular weight
components for total joint replacement. J Bone Joint Surg [Am]
6. Andriacchi PT, Mikosz RP. Musculoskeletal dynamics, locomotion
and clinical applications. In: Mow VC, Hayes WC, ed. Basic Ortho-
paedic Biomechanics. New York, Raven Press, 1991:51-92.
7. Morrison JB. Function of the knee joint in various activities. Biomed
Engin 1969;4:573-80.
8. Morrison JB. The mechanics of the knee joint in relation to normal
walking. J Biomech 1970;3:51-61.
Table I. Reported tibiofemoral joint loads for several daily activities
Tibiofemoral joint load
Author(s) Activity (body-weight multiples)
Ericson and Nisell
Cycling 1.2
Level walking 3.0
Present study 3.9
3.9 to 6.0
Downstairs walking 3.8
Downhill walking 4.5
Present study 8
Rising from chair 3.2
Isokinetic knee extension 9
at 30°/sec
Squat descent 5.6
Weightlifting (120 kg) up to 24
Table II. Reported patellofemoral joint loads for several daily activities (NB:
where the data for the joint load calculations were taken from secondary sources then
these are indicated in parentheses)
Patellofemoral joint load
Author Activity (body-weight multiples)
Ericson et al
Cycling 1.2
Reilly and Martens
Level walking 0.5
) 0.7
) 1.3
) Stair ascending 2.1
Reilly and Martens
) 2.5
) Stair descending 5.6
Reilly and Martens
) 2.5
) Downhill walking 1.8
Rising from chair 3.1
) Jogging 7.0
Squat descent 7.6
Isometric contraction at 6.5
90° flexion
Isokinetic exercise at 12.0
) Kicking 7.6
Patellar tendon rupture 17.5 - 25
9. Collins JJ. The redundant nature of locomotor optimization laws.
J Biomech 1995;28:251-67.
10. Jefferson RJ, Collins JJ, Whittle MW, Radin EL, O’Connor JJ.
The role of the quadriceps in controlling impulsive forces around heel
strike. Proc Inst Mech Eng H 1990;204:21-8.
11. Collier JP, Mayor MB, McNamara JL, Suprenant VA, Jensen RE.
Analysis of the failure of 122 polyethylene inserts from uncemented
tibial knee components. Clin Orthop 1991;273:232-42.
12. Mintz L, Tsao AK, McCrae CR, Stulberg SD, Wright T. The
arthroscopic evaluation and characteristics of severe polyethylene
wear in total knee arthroplasty. Clin Orthop 1991;273:215-22.
13. Kilgus DJ, Moreland JR, Finerman GA, Funahashi TT, Tipton JS.
Catastrophic wear of tibial polyethylene inserts. Clin Orthop
14. Wasielewski RC, Galante JO, Leighty RM, Natarajan RN, Rosen-
berg AG. Wear patterns on retreived polyethylene tibial inserts and
their relationship to technical considerations during total knee arthro-
plasty. Clin Orthop 1994;299:31-43.
15. Winter DA. Biomechanics and motor control of human movement.
2nd ed. New York, John Wiley, 1990.
16. Nisell R. Mechanics of the knee: a study of joint and muscle load with
clinical applications. Acta Orthop Scand 1985;56:[Supp] 216.
17. Campion PJ, Burns JE, Williams A. A code of practise for the
detailed statement of accuracy. National Physics Laboratory Publica-
tion, London, Her Majesty’s Stationery Office, 1973.
18. Kuster M, Sakurai S, Wood G A. Kinematic and kinetic comparison
of downhill and level walking. Clin Biomech 1995;10:79-84.
19. Kuster M, Sakurai S, Wood G A. Downhill walking: a stressful task
for the anterior cruciate ligament? Knee Surg Sports Traumatol
Arthroscopy 1994;2:2-7.
20. Walker PS. Requirements for successful total knee replacements:
design considerations. Orthop Clin North Am 1989;20:15-29.
21. McNamara JL, Collier JP, Mayor MB, Jensen RE. A comparison of
contact pressures in tibial and patellar total knee components before
and after service in vivo. Clin Orthop 1994;299:104-13.
22. Wyss UP, Costigan PA, Li J, Olney SJ, Zee BC, Cooke TDV. Bone-
on-bone forces at the knee joint during walking and stairclimbing.
Proc XIVth Congress of ISB, Paris 1993:1482-3.
23. Collier JP, McNamara JL, Suprenant VA, Jensen RE, Suprenant
HP. All-polyethylene patellar components are not the answer. Clin
Orthop 1991;273:198-203.
24. Buechel FF, Pappas MJ, Makris G. Evaluation of contact stress in
metal-backed patellar replacements: a predictor of survivorship. Clin
Orthop 1991;273:190-7.
25. Kuster M, Wood GA, Sakurai S, Blatter G. Stress on the femor-
opatellar joint in downhill walking: a biomechanical study. Z Unfall-
chir Versicherungsmed 1993;86:178-83.
26. Smith AJ. A study of forces on the body in athletic events with
particular reference to jumping. PhD Thesis, Leeds, England, 1972.
27. Zernicke RF, Garhammer J, Jobe FW. Human patellar-tendon
rupture: a kinetic analysis. J Bone Joint Surg [Am] 1977;59-A:
28. Bartel DL, Rawlison JJ, Burstein AH, Ranawat CS, Flynn WF Jr.
Stresses in polyethylene components of contemporary total knee
replacements. Clin Orthop 1995;317:76-82.
29. Kettelkamp DB, Jacobs AW. Tibiofemoral contact area: determina-
tion and implications. J Bone Joint Surg [Am] 1972;54-A:349-56.
30. Fukubayashi T, Kurosawa H. The contact area and pressure distribu-
tion pattern of the knee. Acta Orthop Scand 1980;51:871-9.
31. Postak PD, Heim CS, Greenwald AS. Tibial plateau surface stress in
TKA: a factor influencing polymer failure. AAOS 1994.
32. Argenson JN, O’Connor JJ. Polyethylene wear in meniscal knee
replacement: a one to nine-year retrieval analysis of the Oxford knee.
J Bone Joint Surg [Br] 1992;74-B:228-32.
33. Plante-Bordeneuve P, Freeman MAR. Tibial high-density polyethyl-
ene wear in conforming tibiofemoral prostheses. J Bone Joint Surg
[Br] 1993;75-B:630-6.
34. Ericson MO, Nisell R. Tibiofemoral joint forces during ergometer
cycling. Am J Sports Med 1986;14:285-90.
35. Harrington IJ. A bioengineering analysis of force actions at the knee
in normal and pathological gait. Biomed Eng 1976;11:167-72.
36. Andriacchi PT, Andersson GBJ, Fermier RW, Stern D, Galante
JO. A study of lower-limb mechanics during stair-climbing. J Bone
Joint Surg [Am] 1980;62-A:749-57.
37, Ellis MI, Seedhom BB, Wright V. Forces in the knee joint whilst
rising from a seated position. J Biomed Eng 1984;6:113-20.
38. Dahlkvist NJ, Mayo P, Seedhom BB. Forces during squatting and
rising from a deep squat. Eng Med 1982;11:69-76.
39. Collins JJ. Antagonistic-synergistic muscle action at the knee during
competitive weight lifting. Med Biol Eng Comput 1994;32:168-74.
40. Ericson MO, Nisell R. Patellofemoral joint forces during ergometric
cycling. Phys Ther 1987;67:1365-69.
41. Reilly DT, Martens M. Experimental analysis of the quadriceps
muscle force and patello-femoral joint reaction force for various
activities. Acta Orthop Scand 1972;43:126-37.
42. Matthews LS, Sonstegard DA, Henke JA. Load bearing character-
istics of the patello-femoral joint. Acta Orthop Scand 1977;48:511-6.
43. Boccardi S, Pedotti A, Rodano R, Santambrogio GC. Evaluation of
muscular moments at the lower limb joints by an on-line processing of
kinematic data and ground reaction. J Biomech 1981;14:35-45.
44. Morrison JB. The mechanics of the knee joint in relation to normal
walking. J Biomech 1970;3:51-61.
45. Winter DA. Moments of force and mechanical power in jogging.
J Biomech 1983;16:91-7.
46. Huberti HH, Hayes WC. Patellofemoral contact pressures: the
influence of Q-angle and tendofemoral contact. J Bone Joint Surg
[Am] 1984;66-A:715-24.
47. Nisell R, Ericson M. Patellar forces during isokinetic knee extension.
Clin Biomech 1992;7:104-8.
48. Wahrenberg H, Lindbeck L, Ekholm J. Knee muscular moment,
tendon tensio force and EMG during a vigorous movement in man.
Scand J Rehabil Med 1978;10:99-106.
VOL. 79-B, N
. 1, JANUARY 1997
... However, although included in most recommendations for the treatment of OA and addressed as a modifiable risk factor [39], weight loss does not occur frequently in practice. For every kilogram of body weight lost, the knee experiences a fourfold reduction in load during daily physical activity [40]. Regardless of obesity, whether hyperandrogenism, low-grade inflammation, and metabolic changes may further increase the risk of OA (Figure 1) remains a question that warrants further research. ...
Full-text available
The close link between osteoarthritis (OA) and metabolic disorders on the one hand and hormonal disorders on the other suggests a possible association between OA and endocrine-metabolic disorders, such as PCOS. The aim of this review is to analyze the relationship between PCOS and OA, to consider the common pathogenetic mechanisms between the two conditions, and to summarize the data accumulated so far in the literature. For the purposes of our narrative review, a comprehensive search was conducted within credible databases. Our literature search found that epidemiological studies have shown a higher incidence of knee and hip OA in women with PCOS. This can be partly explained by obesity, which is a common intersection between the two conditions. Potential mechanisms among OA, PCOS, and obesity were considered. Another common point between OA and PCOS is that both conditions can be considered as highly heterogeneous syndromes with various etiologies, the result of a combination of systemic (genetic, hormonal, and metabolic) and local factors. To date, hyperandrogenism and greater cartilage thickness in young women with PCOS remain unclear in terms of determining the risk of developing OA. Prospective longitudinal studies are needed to assess the “fate” of the weight-bearing joints in women with PCOS, who are more likely to suffer from knee joint complaints.
The knee is the largest and one of the most biomechanically demanded joints in the human body, as it is located between the two longest lever arms of the body and its most powerful muscles. Knee stability through the range of motion is ensured by both static and dynamic structures that work in concert to prevent excessive movement or instability, which may occur across multiple planes of motion. While offering a wide range of motions, these structures have to meticulously balance the compressive forces across the knee joint. Orientation, shape, and material properties of the bony structure and dynamic stabilizers, including ligaments, the capsule, and musculotendinous soft tissues, are essential for knee stability. Even small changes to any of these parameters will alter the inherently complex interactions between these structures and ultimately distort overall movement patterns of the knee, consequently impacting alignment, load distribution, and wear of the components forming the knee joint.
Background Biomechanical study is fundamental for the preclinical evaluation of knee prostheses. However, there are few reports on the contact characteristic investigation in the hinged knee prosthesis. The purpose of this study was to investigate the contact characteristics of a novel hinged knee prosthesis. Methods All of the component models were designed and assembled using Solidworks. A comparison of the contact area and ultra-high-molecular-weight polyethylene (UHMWPE) deformation using the experimental method (EM) and finite-element analysis (FEA) under 3000 N with the prosthesis at different flexions was performed. The peak contact pressure and von Mises stress on tibial insert and bushing were investigated under nine specific samples that were extracted from a gait cycle using FEA (according to ISO 14243-1: 2009). Results The largest contact area and UHMWPE deformation were 100.78 ± 8.71 mm² and 0.085 ± 0.015 mm in the EM, and 96.68 mm² and 0.096 mm in FEA. The peak contact pressure and von Mises stress on the tibial insert were 26.3071 MPa and 10.5115 MPa at 13% of the gait cycle and on bushing were consistently 0 MPa. The contact pressures were distributed at the posterior of the insert. Conclusion The finite-element model was validated to be applicable for predicting the real prosthesis behavior based on the good correlation of the results using the EM and FEA. The model can help to identify contact characteristics and be can used in optimization studies of this novel prosthesis during the design phase.
Full-text available
Objectives: To investigate the contact stress and the contact area o tibial inserts and bushings with respect to different congruency designs in a spherical center axis and rotating bearing hinge knee prosthesis under gait cycle loading conditions using finite element analysis. Methods: Nine prostheses with different congruency (different degrees of tibiofemoral conformity and different distances between the spherical center and the bushing) designs were developed with the same femoral and tibial components. The models were transferred to finite element software. The peak contact stresses and contact areas on tibial inserts and bushings under the gait cycle loading conditions were investigated and compared. Results: For tibial insert, the peak contact stress was the highest in the low conformity-long group (61.4486 MPa), and it was 1.88 times higher than that in the group with the lowest stress (moderate conformity-short group, 32.754 MPa). The contact area was the largest in the low conformity-long group (420.485 mm2 ), and it was 1.19 times larger than that in the group with the smallest area (moderate conformity-middle group, 352.332 mm2 ). For bushing, the peak contact stress was the highest in the high conformity-long group (72.8093 MPa), and it was 3.21 times higher than that in the group with the lowest stress (high conformity-short group, 22.6928 MPa). The contact area was the largest in the low conformity-short group (2.41 mm2 ), and it was 2.27 times larger than that in the group with the smallest area (high conformity-middle group, 1.063 mm2 ). Conclusion: The results of our study showed that the congruency of the tibiofemoral surface and bushing surface should be considered carefully in the design of the spherical center axis and rotating bearing hinge knee prosthesis. Different levels of contact performance were observed with different congruency designs. In addition, the influence of contact stress and contact area on the polyethylene wear of rotating hinge knee prostheses should be confirmed with additional laboratory tests.
Objectives: To investigate the effect of screw length, lateral hinge fracture, and gap filling on stability after medial opening wedge high tibial osteotomy (MOW HTO) using a locking plate. Methods: Forty tibiae from fresh-frozen cadavers were randomly allocated into five groups. Group A was bicortically fixated, while Group B and Group C were unicortically fixated: 90% and 55% of drilled tunnel length, respectively. Group D was fixated using 90% length screws with a fractured lateral hinge. Group E was fixated using 90% length screws with gap filling using a bone substitute. Operated tibiae were tested under axial compressive load using a material testing machine. The medial gap changes under the serial axial load of 100-600 N and ultimate failure load were measured. Results: Group D showed the biggest medial gap change and lowest failure load, while Group E presented the smallest gap change and highest failure load. The medial gap changes tended to increase with shorter screw length, but the difference was not significant between Groups A, B, and C. Group C and Group D showed greater medial gap change and lower failure load compared with Group E, while not differing from Group A and Group B. Conclusions: Unicortical fixation in proximal screw holes of a locking plate was not inferior to bicortical fixation regarding axial stability in MOW HTO, although proximal screws that are too short should be avoided. Lateral hinge fracture decreased, while gap filling with bone substitute increased axial stability.
Full-text available
Knee arthrodesis is commonly performed to treat joint pain and instability in cases of multiple failed knee replacement surgeries, bone loss, traumatic injury, infection, or loss of the quadriceps extensor mechanism, but its mechanics are poorly understood. This study quantified the changes in gait kinematics and kinetics induced by simulating knee arthrodesis. A total of 10 healthy subjects (M/F, n = 5/5; aged 18–23) with no history of gait abnormalities were recruited for this study, and knee arthrodesis was simulated for 0 through 24 h using an immobilizing knee brace. Gait analysis was conducted during level walking on a 15-m walkway using an eight-camera motion tracking system with a six-axis ground reaction force platform. Normal gait served as each subject’s control, and measurements of braced gait were taken immediately following the fitting of the brace and over a 24-h acclimation period. Results showed that simulated knee arthrodesis caused statistically significant and interrelated changes to gait when compared to normal gait, including reduced ankle plantar flexion and pelvic obliquity during swing phase of the braced limb, increased abduction of the left and braced hip during stance of the braced limb, and an increase in the net joint reaction force at the contralateral hip. These changes facilitated ground clearance of the braced limb, but however did not show increases in knee flexion moment that might serve as a cyclic fatiguing factor in implants or fusions used to fixate the knee. Increased acclimation time was not found to significantly alter braced gait kinematics, supporting the use of short-term bracing as an experimental model for the longer-term immobilized knee.
Full-text available
Optimal management of knee osteoarthritis (KOA) should include, where possible, modification of risk factors through targeted interventions. The objectives of the present narrative review were to identify, summarize, and cluster all the potentially modifiable risk factors that influence the course of KOA, and discuss their susceptibility to alteration via personal, clinical, and public strategy. For this purpose, Pubmed and Scopus databases were queried using the terms “knee osteoarthritis”, “risk factors” and “improvement”. Six main categories of modifiable risk factors were identified: (1) obesity and overweight, (2) comorbidity, (3) occupational factors, (4) physical activity, (5) biomechanical factors, (6) dietary exposures. In the era of age- and obesity-related diseases, the combined effects of local and systemic risk factors should be managed by combined measures. Femoral muscle-strengthening physical activities, complemented with proper diet, weight loss, vocational rehabilitation, management of comorbidities (especially diabetes and depression), and biomechanical support may add up to the holistic therapeutic approach towards the patient with KOA. An individual risk factor modification program should be developed in accordance with patient preferences and habits, workplace, medical history, and overall health condition. Due to its great impact on a wide range of functions and tissues, interventions on modifiable risk factors improve not only the symptoms of KOA but also affect the osteoarthritic joint as a whole.
During in vivo function, implant surfaces are exposed to a corrosive environment as well as a wide range of tensile and compressive stresses. The objective of this study was to study the corrosion behavior of the Ti–6Al–4V orthopaedic alloy under different levels (ranging from plastic to elastic) of tensile stress, compressive stress, and combined tensile and compressive stresses in phosphate buffered saline 1X solution. The introduction of stress was found to alter the corrosion behavior of the Ti–6Al–4V alloy. The corrosion behavior was slightly improved under compressive stress while specimens under tensile stress showed weaker corrosion resistance compared to the specimens under zero stress. Nonetheless, the synergistic action of the tensile and compressive stresses considerably increased the corrosion activity of the specimens, more than that when they were just under tensile stress. The worst corrosion behavior was observed for the specimens under coupled elastic stress.
Conference Paper
The purpose of this study was to investigate the effect of changing the plantar flexion resistance (PFR) of an ankle-foot orthosis (AFO) on the compressive tibiofemoral force, knee muscle forces, and knee joint angle. We measured and estimated knee flexion angle, knee muscle force, and the compressive tibiofemoral force in healthy adult males. The results showed that the first peak compressive tibiofemoral force, peak knee flexion angle, and peak quadriceps muscle force increased in the strong PFR condition compared with the no-AFO condition. These results suggest that over-PFR caused various knee troubles.
The success of joint replacement surgery has been responsible for raising patients' expectations regarding the procedure. Many of these procedures are currently designed not only to relive the pain caused by arthrosis, but also to enable patients to achieve functional recovery and to engage in some degree of physical activity and sports. However, as physical exercise causes an increase in forces exercised through the articular prosthesis, it can be an important risk factor for its early failure. Scientific literature on sports after arthroplasty is limited to small-scale retrospective studies with short-term follow-up, which are mostly insufficient to evaluate articular prosthesis durability. This article presents a review of the literature on sports in the context of hip, knee, shoulder and intervertebral disc arthroplasty, and puts forward general recommendations based on the current scientific evidence. Systematic Review, Level of Evidence III.
Eight healthy male subjects performed isokinetic maximum knee extensions from 90 degrees flexion to full extension in a CYBEX n apparatus at two different speeds (30° and 180° s(-1)). Using a planar biomechanical model of the patellofemoral joint, the patellar forces in the sagittal plane were quantified. At the slower speed the patellofemoral compressive force and the suprapatellar tendon force reached values of about 12 bodyweights while the infrapatellar tendon force did not exceed 9 bodyweights. At the faster speed, the corresponding force magnitudes were 7.5 bodyweights and 5.5 bodyweights. The force peaks occurred at the beginning of the extension movement between 65° and 75° of knee flexion and were a function of knee angle and knee extension strength. The magnitude of the patellar forces during isokinetic knee extension of maximum effort were compared to other knee extending activities and were found to be considerably higher than during walking, jogging, and cycling.
The classic book on human movement in biomechanics, newly updated. Widely used and referenced, David Winter's Biomechanics and Motor Control of Human Movement is a classic examination of techniques used to measure and analyze all body movements as mechanical systems, including such everyday movements as walking. It fills the gap in human movement science area where modern science and technology are integrated with anatomy, muscle physiology, and electromyography to assess and understand human movement. In light of the explosive growth of the field, this new edition updates and enhances the text with: Expanded coverage of 3D kinematics and kinetics. New materials on biomechanical movement synergies and signal processing, including auto and cross correlation, frequency analysis, analog and digital filtering, and ensemble averaging techniques. Presentation of a wide spectrum of measurement and analysis techniques. Updates to all existing chapters. Basic physical and physiological principles in capsule form for quick reference. An essential resource for researchers and student in kinesiology, bioengineering (rehabilitation engineering), physical education, ergonomics, and physical and occupational therapy, this text will also provide valuable to professionals in orthopedics, muscle physiology, and rehabilitation medicine. In response to many requests, the extensive numerical tables contained in Appendix A: "Kinematic, Kinetic, and Energy Data" can also be found at the following Web site:
This study of normal patello-femoral biomechanics defines some functional specifications which may be useful in future total knee prosthesis design. Serial lateral X-rays of 15 fresh knees and their patellar mechanisms at several flexion angles provided definition of the direction of the resolved patello-femoral forces. Assuming that the patella acts as a frictionless pulley, the magnitude of the patello-femoral forces during several activities was calculated using data from Morrison (1970) and Smidt (1973). It ranged between 421 and 3420 newtons for the various activities and for isometric exercise. A methylene blue contact print technique was used to measure the bearing areas. These data indicate that between 13 and 38 per cent of the patellar surface bears joint loadings. Patello-femoral contact stresses were calculated to range from 1.28 to 12.6 N/mm2. A 696 new-ton man climbing stairs would, for example, generate a patello-femoral force of 1754 newtons and would experience patello-femoral contact stresses between 3.73 and 6.87 N/mm2. Stress values were equal to or in excess of anticipated tibial-femoral stresses. The high patello-femoral load values, the small bearing surfaces, and the consequent significant stress magnitudes indicate the need for caution in development of a patello-femoral joint prosthetic replacement.
With injuries to the components of the extensor apparatus of the knee as a background, it is interesting to investigate the magnitude of forces acting on these components, i.e. m. quadriceps femoris, the quadriceps tendon, patella, lig. patellae and tuberositas tibiae, during a vigorous but physiological movement. By means of the dynamic laws of mechanics the muscular moment of force with respect to the bilateral knee axis during kicking was calculated in 6 normal subjects. It was found that the maximum extending muscular moment in the knee occurs very early in the movement, when the initial flexion changes into extension, and thus long before the ball is hit. The peak of quadriceps EMG activity coincides with maximum moment. The EMG peak of the antagonistically acting hamstrings comes later, nearer to when the ball is struck. The greatest extending muscular moment obtained during the swing phase of kicking was surprisingly high, 260 Nm, corresponding to a tension force in the patellar tendon of 5200 N or about 7 times body weight. These values are discussed in relation to tendon strength.
The first biomechanical analysis of a human patellar-tendon rupture during actual sports competition is reported. Cinematographic data for analysis were collected at a national weight-lifting championship. Dynamic equations to mathematically model the lifter were developed to compute time course and magnitudes of hip, knee and ankle-joint moments of force and of tensile loading of the patellar tendon before and during tendon trauma. Results provided evidence that the range of maximum tensile stress of the tendon may be considerably greater during rapid dynamic loading conditions, as in many sports situations, than maximum tensile stress obtained during static test conditions.
The paper describes a simplified bioengineering analysis for the determination of force actions at the knee joint for normal and pathological gaits. The stance phase of the gait cycle only is considered, and gravitational and inertia forces are excluded from the analysis. The anatomical and functional assumptions required for analysis are discussed. Force actions transmitted at the knee joint by the bearing surfaces, muscles and ligaments for normal individuals are presented with reference to magnitude and phasic relationship for the activity of level walking. Bearing loads transmitted at the knee joints of pathological limbs are also discussed.
We recovered 23 meniscal bearings from 18 failed bicompartmental Oxford knee prostheses. They had been implanted for one to nine years. The minimum thickness of the retrieved bearings was measured and compared with the thickness of 25 unused bearings. The mean penetration rate, calculated by two methods, was either 0.043 or 0.026 mm per annum. This compares with 0.19 mm per annum reported for the Charnley hip. The use of a fully congruous meniscal bearing prosthesis can reduce wear in knee arthroplasty to a very low rate.
We examined 86 polyethylene inserts, retrieved from total and unicompartmental knee prostheses after an average of 39.5 months in situ, grading them from 0 to 3 for seven modes of polyethylene degradation. Severe wear, with delamination or deformation, was observed in 51% of the implants, and was associated with time in situ, lack of congruency, thin polyethylene, third-body wear debris, and heat-pressed polyethylene. Significant under-surface cold flow was identified in some areas of unsupported polyethylene, and was associated with delamination in the load-bearing areas of thin inserts above screw holes in the underlying metal tray. We recommend the use of thicker polyethylene inserts, particularly in young, active patients and in designs with screw holes in the tibial baseplate. Thin polyethylene inserts which are at risk for accelerated wear and premature failure should be monitored radiographically at annual intervals.