U.S. patent number 3,909,855 [Application Number 05/523,375] was granted by the patent office on 1975-10-07 for below-the-knee prosthesis.
Invention is credited to Joseph G. Barredo.
United States Patent |
3,909,855 |
Barredo |
October 7, 1975 |
Below-the-knee prosthesis
Abstract
A hollow, rigid, lightweight nonarticulated prosthesis with a
foreshortened foot fits the stump of a below-the-knee amputee and
has an optional simple above the knee-cap holding strap to secure
the prosthesis when the knee is straightened. The center of mass of
the prosthesis is much closer to the knee than is the center of
mass of a natural leg.
Inventors: |
Barredo; Joseph G. (Washington,
DC) |
Family
ID: |
24084737 |
Appl.
No.: |
05/523,375 |
Filed: |
November 12, 1974 |
Current U.S.
Class: |
623/32; 623/27;
623/33 |
Current CPC
Class: |
A61F
2/60 (20130101); A61F 2002/5001 (20130101) |
Current International
Class: |
A61F
2/60 (20060101); A61F 001/08 (); A61F 001/02 () |
Field of
Search: |
;3/2,6,7,16-20 |
References Cited
[Referenced By]
U.S. Patent Documents
Primary Examiner: Frinks; Ronald L.
Attorney, Agent or Firm: Wolfe, Hubbard, Leydig, Voit &
Osann, Ltd.
Claims
Wherefore, I claim as my invention:
1. A walking prosthesis for fitting a below-the-knee stump
comprising a hollow, one-piece, rigid, lightweight leg
corresponding generally in form to the human leg having a socket
fitted to receive the stump, and with a weight several times less
than that of the natural leg and the foot portion shortened at the
toe end by an amount approximating the natural toe length.
2. The prosthesis of claim 1 in which the bottom of the foot
portion curves upward at the toe end to provide a rocking
surface.
3. The prosthesis of claim 1 with a strap loop secured at its ends
to the sides of the hollow leg near its top and with a length short
enough to engage the wearer's leg above the knee cap when the knee
is straightened for holding the prosthesis in place.
4. The prosthesis of claim 1 with knee-hinge engaging facing
concave side extensions to grasp the ends of the knee-hinge for
holding the prosthesis in place.
Description
This invention relates to an improved below-the-knee
prosthesis.
An artificial leg for a person such as myself who has a
below-the-knee amputation and has tissues sensitive to the
pressures and friction associated with a use of a prosthesis
presents a particularly vexing problem if a reasonably full use of
the knee is to be preserved. A prosthesis can usually be fitted to
the stump of the leg below the knee and strapped to the leg above
the knee so that use is made of the knee in walking. With approved
prostheses now available, this typically results in much more
expenditure of energy for walking at the same comfortable speed
that the user would enjoy with natural limbs. If demands on the
heart of the conventional prosthesis wearer do not prevent his
escape from invalidism, fatigue and discomfort are likely from
extended use, and pathological conditions--cysts, blisters, edema,
and infections--may result from friction or pressure by the
artificial leg and harness against the skin and other tissues.
Moreover, when peripheral vascular diseases lead to amputation of
the lower extremity, it is well known that the more distal the site
of the amputation the greater the incidence of poor healing, while
the more proximal the site of amputation, the better the potential
for healing but the higher will be the expenditure of energy in
walking and the greater will be the degree of invalidism. Thus, the
surgeon and his patient face the difficult choice of saving length
at the risk of impaired wound-healing, or sacrificing length at the
risk of impaired function. No available prosthesis was heretofore
designed to eliminate this hard choice.
The direction of development in artificial legs has generally been
from the simple to the complex, with more or less articulated toe
and ankle action in an attempt to initiate nature by duplicating
the functions of the natural foot. The importance of the cosmetic
aspects is not to be denied, beyond a certain point simulating the
joints and weight of the natural leg is self-defeating a certain
point. The prosthesis is not a natural limb and cannot function in
all respects as a natural, integral limb. Disfiguring or
distracting impediments in the user's gait and carriage are a
significant cosmetic fault likely to accompany the relatively
heavy, jointed prosthesis. Problems have also been introduced in
attempts to simplify the prosthesis. A currently favored
simplification is the SACH (Solid Ankle, Cushioned Heel) foot
designed to be fixed to the prosthesis. The SACH ankle, being
solid, does not pivot or hinge to change the attitude of the foot
as the heel engages the surface being walked upon, Instead the
resilient cushioned heel yields to absorb the energy of the impact
of the mass of the user's body. To some extent, the impact energy
stored in natural walking speed without additional energy
expenditure in an elastic heel facilitates pushing the heel end of
the prosthesis upward as weight is transferred to the other leg in
walking, depending upon the walking step and speed. The energy
required to accelerate or decelerate the mass of the prosthesis in
walking is by no means negligible, but less appreciated is the fact
that the wearer of the prosthesis must stabilize his body to
counteract the forces of deacceleration as well as acceleration,
and the tissues of the stump or the knee of the fitted leg may be
unduly stressed. Elaborate or close fitting harnesses must often be
employed to better accommodate the control of the prosthesis and
distribute the reaction forces, as well as to keep the prosthesis
from flying off during vigorous movements.
It is the purpose of my invention to provide a below-the-knee
prosthesis which permits the user to walk at a natural walking
speed without additional energy expenditure than in walking with
natural limbs and which minimizes discomfort and other ill effects
suffered by users of conventional prostheses.
It is likewise a principal object of my invention to provide a new
type of below-the-knee prosthesis that accommodates a proximal site
of amputation without increasing the expenditure of energy in
walking or the degree of invalidism. It permits walking with the
same comfortable walking speed (CWS) as a normal person.
My new prosthesis eliminates the excess expenditure of energy with
existing prostheses which, besides loading the heart, produces such
pathological conditions as blisters, edema, and infections. My new
type of prosthesis avoids the need of eliminating the knee joint
when the length of the stump would be too short to walk with the
existing type of prosthesis.
I have approached the design of my prosthesis by recognizing that
(1) the mechanics of walking with an artificial leg be considered
from an entirely different viewpoint than walking with a natural
leg, and (2) that fatigue and possible pathological events
associated with the harness on the above-the-knee tissue supporting
the below-the-knee prosthesis follows from the forces involved in
the mechanics of walking with the prosthesis. The key to
accommodating both lies in minimizing the prosthesis acceleration
and deceleration forces encountered in walking at normal speeds to
thus minimize the prosthesis reaction forces against the supporting
tissues. The lesser energy expenditure lessens the load on the
heart involved of the cardiac patient whose use of a conventional
prosthesis is otherwise severely restricted.
As a rule of thumb, this approach calls for a high location of
center of mass of the combined stump and prosthesis so that the
distance between the hinge of the knee and the center of mass is
small. The prosthesis itself must be structured (a) to be very
light and (b) to have itself a relatively high (i.e., towards the
knee end) center of mass. As far as my design is concerned. any
theory that the weight and weight distribution of the prosthesis
should approximate that of the limb it replaces is completely
wrong.
To further control acceleration forces so as to limit friction and
pressures on the tissues of the prosthesis user, I have
foreshortened the foot portion to eliminate that length
corresponding to the toes of the natural foot and have made the
prosthesis rigid with neither an ankle joint nor cushion heel.
Following these principles permits elimination of the prosthesis
harness, or if the user feels more secure if a harness is used, the
use of a vastly simplified harness, which may consist of a simple
loop attached to the upper sides of the prosthesis for riding the
leg tissue above the knee cap. FAtigue and pathological problems
associated with extended reliance upon a prosthesis are minimized
and with a desired simplicity, economy, and preservation of
cosmetic values.
Other advantages and objects of my invention will become apparent
in the following detailed description of a preferred embodiment of
my invention as read in connection with the accompanying drawings,
in which:
FIG. 1 is a perspective view, partially cut away, of a strapless
below-the-knee prosthesis embodying my invention;
FIG. 2 is a side view of the prosthesis of FIG. 1 illustrating how
it fits the knee joint and also showing the foreshortened foot
portion of the prosthesis with a lightweight shoe in dotted
outline;
FIG. 3 is a front view of the strapless prosthesis of FIG. 1
partially in action to show the socket for the stump;
FIG. 4 is a side view of the upper portion of a prosthesis
generally similar to that of FIG. 1 but incorporating a simple
strap;
FIG. 5 shows the FIG. 4 prosthesis with harness strap on the leg of
a user whose knee is bent as during a sitting position;
FIG. 6 is a series of leg and hip schematic sketches showing, from
right to left, successive stages in a walking step.
Turning now to FIGS. 1, 2, and 3, a below-the-knee prosthesis for
an artificial leg 10 is shown in the form of a unitary hollow shell
preferably formed of fiberglass over a mold. The outer surface of
the fiberglass is preferably smoothly finished and colored for the
desired cosmetic effect. The prosthesis wall thickness may be
conveniently small, for example, on the order of one-eighth inch or
even smaller, and yet be both strong and rigid without resort to
internal bracing. As a result, the prosthesis is several times
lighter than the natural leg. which is one of the design
objectives. In my own case, as an adult male weighing 160 pounds
with a short below-the-knee stump as in FIG. 2, I am accustomed to
a normal walking rate and to going up and down stairs with a
fiberglass prosthesis embodying my invention which weighs only 19
ounces. This has proven more than adequate in strength, durability,
and comfort. I do not wish to suggest that 19 ounces is or should
be a lower weight limit, and the economic availability of other
materials and manufacturing processes may make lighter weights
practical. The prosthesis is significantly not only several times
lighter than the natural leg but also several times lighter than
the usually prescribed prosthesis with resulting beneficial
decrease of accelerating and decelerating forces during
walking.
For external fitting, the prosthesis has concave side extensions or
caps 11 at the top of the prosthesis which extend above the
cut-down front portion 12 and back portion 13 to fit closely over
the ends of the knee hinge. The side portions 11 are sufficiently
resilient to grip the sides of the knee sufficiently to hold the
prosthesis in place. The front portion 12 of the prosthesis must be
cut away enough so that it does not bear against the patella and
the rear portion is cut away enough to avoid the top of the
prosthesis bearing against the upper leg tissues when the knee is
bent.
A smooth surfaced socket 14, also suitably molded of fiberglass, is
secured to the upper end of the prosthesis to receive the stump.
The stump socket 14 is preferably custom molded as by making a
plaster replica of the stump as the form for the socket. The socket
is then cemented in place in the prosthesis. The socket is designed
for a snug fit but does not clamp the stump. The usual precautions
for controlling perspiration problems are followed, it being common
to stretch a knit stocking over the stump before inserting it in
the socket or to provide vents in the socket itself.
The shape and rigidity of the foot portion 15 of the prosthesis,
and not only its lightness, are of considerable importance for ease
of walking and comfort, Overall, it is a lightweight, rigid shell
and an integral part of the prosthesis. It may be made lighter by
apertures (not shown) in the arch of the prosthesis or elsewhere
not distractingly noticeable in the foot portion. Very
significantly, the foot portion is also foreshortened at the toe
end 16 by a length approximately equal to the toes of the natural
foot. As a cosmetic necessity, a lightweight shoe 17 (FIG. 2),
matching the shoe on the user's other natural foot, is worn over
the prosthesis. While any shoe weight at all offsets in part the
desirable weight reduction and weight distribution achieved by the
construction of the prosthesis, this only emphasizes the importance
of weight control in the prosthesis itself in order that the shoe
weight can be better tolerated. In my own case I have found an
inexpensive 10 -ounce medium size men's shoe to be quite
satisfactory.
The shoe on the prosthesis should be sufficiently flexible (as is
normally the case in an inexpensive, lightweight shoe) to bend
easily at the toe end when worn over the prosthesis during walking.
The toe end of the foot 15 is not only shortened, but the bottom of
the front portion 18 of the foot desirably curves upward rather
sharply as shown within the normally dimensioned confines of the
shoe. This shape defines a curved fulcrum about which the foot may
rock as the angle of the prosthesis foot with respect to the ground
changes during walking.
The wearer of the prosthesis, whether or not he is wearing a shoe
over the prosthesis, benefits in walking on the foreshortened
length of the prosthesis foot. The foreshortened prosthesis foot
avoids the hip elevation and additional acceleration forces which
would be encountered in walking with a longer foot. The lightweight
shoe 17 has no provision for large heel resilience, nor is any
sought since the acceleration caused by the release of energy of a
very resilient cushion is more tiring to the prosthesis wearer than
any shock occurred when the substantially uncushioned rigid heel
contacts the ground during walking.
As may be noted further in FIG. 2, the arch 19 of the foot portion
of the prosthesis is preferably greater in height than that of the
normal foot, the smaller radius of curvature lending itself to a
stronger structure for a given amount of material.
A slightly different version of the prosthesis is shown in FIGS. 4
and 5 where a strap is provided to hold the prosthesis in place
against the forces of gravity and other accelerations without
reliance upon a close fit between the upper side extensions of the
prosthesis and the ends of the knee hinge. This may be advantageous
where the knee hinge convexity is not pronounced, where the user
would rather avoid close contact between the prosthesis and the
knee hinge, or simply where the user feels insecure without the
more assurance provided by this strap. The strap 20 shown in FIGS.
4 and 5 is suitably a plastic covered wire or other smooth surfaced
cord having its ends inserted in apertures 21 at the sides of the
prosthesis near its top. In this model the upper portions may curve
outward from the knee hinge rather than contact it, and may be less
in height so as to terminate below the hinges. For convenience the
cord ends are inserted through the apertures from an inside surface
of the leg and then knotted outside the prosthesis to maintain a
support loop of the desired length. As shown further in FIG. 4, the
cord 20 must be long enough to pass around the front of the leg
just above the knee cap, the knee cap preventing the cord from
slipping down. The risk of irritation due to the force of the cord
against the leg tissues has thus far proven surprisingly small. The
cord need not be under any appreciable tension; when the wearer is
seated with the knee bent, the cord may be so comfortably loose as
to be lifted from the leg as shown in FIG. 5.
The elimination of the SACH foot and the light weight of the
prosthesis as a whole work in a helpful direction reducing both the
prosthesis mass and its center of mass distance from the knee. The
relatively short distance d from the center of mass of a short
stump to the knee hinge is assompanied by the relatively short
distance D from the center of mass of the prosthesis to the knee
hinge (d and D are indicated in FIG. 2). A lesser rather than
greater amount of energy need be transmitted between the stump and
the prosthesis to provide the sometimes surprisingly high
accelerating and decelerating forces involved in normal walking. A
long stump and a relatively short prosthesis may result in higher
mass and radius values than would a short stump and longer
lightweight prosthesis, although I am unaware of any appreciation
of this concept in prosthesis design prior to my invention.
In the design of the prosthesis of my invention I have arrived at
the low mass and high center of mass by consideration of the basic
physics-- the forces, masses, and motions--of walking. Fitting and
alignment, which usually receive a great deal of attention in the
use of conventional prostheses, are essentially variables of
secondary importance made unnecessarily prominent by slighting
consideration of the primary physical variables. Consider, for
example, that the shorter the walking leg swing time T, the
obviously greater the maximum acceleration and deceleration forces
which must be accommodated. The stump functions at least in part as
a lever for lifting, kicking, or stopping the motion of the
artificial leg. The ratio D/d, although minimized in the design of
my prosthesis by a relatively short D, illustrates the
multiplication of the forces on the stump tissues. The desirably
small amplitude of the prosthesis mass m, which operates on or is
operated by the lever, must be taken into effect. The forces called
for by the feedback signals sent to the brain by the neurovascular
system during walking are largely controlled by these factors. The
measure of these forces can be approximated in part by
consideration of the relationship m 4.pi..sup.2 D.sup.3 /T.sup.2 d.
The human brain considers a great number of relationships and
factors in reaching an end result such as the commodation of time T
to the natural gait of the walker. The brain as a computer takes
into account, for example, the different dimensions in walking
barefoot and walking with heavy shoes. The design of the prosthesis
by considering the physics elements, gives the human brain a
greater range of freedom or effectiveness in realizing a natural
gait within reasonable limits of fatigue and tissue stress.
As further shown in hip and leg sketches of FIG. 6, the prosthesis
also operates to reduce unnecessary vertical displacements of the
body with resulting greater ease in walking. Thus, as shown in the
righthand figure of FIG. 6, the user is kicking his leg to begin a
step. In this part of walking, the prosthesis is being thrown
forward at an appreciable angular velocity around the knee hinge.
Acceleration of the prosthesis causes reaction forces on the leg
stump which can be comfortably borne since the center of mass of
the prosthesis is fairly near that of the stump and because the
mass itself is not high. The deceleration of the leg at the end of
the kick by which the leg is thrown forward in walking again causes
a reaction force of the prosthesis or the prosthesis strap against
the leg, and again the low mass of the prosthesis keeps this force
within comfortable limits. The next stick figure from the right in
FIG. 6 illustrates the following engagement of the prosthesis foot
with the ground after the kick. Contact is made with the heel, and
the foot rocks forward over its curved front portion (the
foreshortened foot length F is indicated in FIG. 2; the FIG. 6 foot
diagrams have no useful scale). The heel does not yield as in a
SACH foot and the user's center of gravity (the hip hinge) changes
only slightly. In the third sketch from the right in FIG. 6, the
prosthesis is not accelerated forward nor is the knee bent as much
as it would otherwise be at the end of the step due to the
foreshortened foot. In the fourth sketch from the right in FIG. 6,
the leg begins to swing from the thigh for the next step to begin.
The knee is not bent as much as it would have been because of the
short foot, and the energy of the next kick about the knee axis
will not be as high.
The fairly constant level of the hip, which the short foot assists,
accommodates walking with the prosthesis with a less awkward and
more natural gait. The gait is not perfectly natural because the
leg is not natural. What is important is that a new accommodation
has been made which minimizes both energy expenditure and
transmittal forces on the leg tissue.
The principles of my invention may be applied to other prostheses
on a moving extremity but it is necessary that the problem be
recognized if the prosthetic solution is to be helpful as in the
foregoing illustration of a below-the-knee example.
* * * * *