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Joint Load Considerations In Total Knee Replacement

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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.
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VOL. 79-B, N
O
. 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,
Switzerland.
G. A. Wood, PhD, Associate Professor
Neuromuscular Performance Laboratory, Department of Human
Movement
G. W. Stachowiak, PhD, Associated Professor
Tribology Laboratory, Department of Mechanical and Materials
Engineering
The University of Western Australia, Nedlands, Australia 6907,
Australia.
Correspondence should be sent to Dr M. S. Kuster.
©1997 British Editorial Society of Bone and Joint Surgery
0301-620X/97/16978 $2.00
JOINT LOAD CONSIDERATIONS IN
TOTAL KNEE REPLACEMENT
MARKUS S. KUSTER, GRAEME A. WOOD, GWIDON W. STACHOWIAK, ANDR
´
E G
¨
ACHTER
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
2
. 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.
1-4
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.
1,5
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.
6
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.
7
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.
9,10
This
finding is consistent with the increasing incidence of reports
of severe wear in joint replacements.
1,2,4,11-14
We present
quantitative joint load data and suggest new criteria for use
in the biomechanical evaluation of total knee prostheses.
MATERIALS AND METHODS
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.
15
A knee joint model, previously described by Nisell,
16
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.
17
Full details of the instrumentation
used can be found in the authors’ previously published
work.
18,19
RESULTS
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
110 M. S. KUSTER, G. A. WOOD, G. W. STACHOWIAK, A. G
¨
ACHTER
THE JOURNAL OF BONE AND JOINT SURGERY
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
2
is
required.
DISCUSSION
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.
19
The
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.
11,20,21
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
9
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
10
found that
the maximum tibiofemoral loads are up to 6.3 times BW,
while Wyss et al
22
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.
23
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.
24
Recent
research indicates higher loadings; for level walking patello-
femoral joint forces of 1.3 to 1.8 times BW,
16,25
for down-
stairs walking 5.5 times BW
16
and for downhill walking 5 to
7 times BW have been suggested.
25
For some sports activ-
ities such as jumping
26
or weightlifting
27
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.
5,11,21,23,24,28
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
2
to 1150 mm
2
.
30
After complete
medial and lateral meniscectomy the tibiofemoral contact
area is approximately 520 mm
2
, depending on the load.
29,30
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
2
depending
on the load, flexion angle and design,
31
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
2
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.
32,33
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
avoided.
111JOINT LOAD CONSIDERATIONS IN TOTAL KNEE REPLACEMENT
VOL. 79-B, N
O
. 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-
theses
31
(TK), a knee joint after meniscectomy
30
(MK) and a natural knee
joint
30
(NK) are shown. The horizontal bar indicates the yield range of
polyethylene.
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
article.
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THE JOURNAL OF BONE AND JOINT SURGERY
Table I. Reported tibiofemoral joint loads for several daily activities
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113JOINT LOAD CONSIDERATIONS IN TOTAL KNEE REPLACEMENT
VOL. 79-B, N
O
. 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. ...
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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.
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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.
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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.