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A Comparative Evaluation of New and Conventional Knee Orthoses for Control of Anterior Tibial Displacement

Glynn, Daniel W. CO; Kennedy, Francis E. PhD; Hood, Matthew O. MS; Greenwald, Richard M. PhD

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JPO Journal of Prosthetics and Orthotics: September 2002 - Volume 14 - Issue 3 - p 113-120
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Although many orthotics and prosthetics facilities still fabricate their own orthoses, the majority of knee orthoses (KO), particularly those used for anterior cruciate ligament (ACL) deficient knees, are purchased from knee brace companies. The sidebars and joints used in commercially available KO are generally not for sale to practitioners, and the cost of tooling that practitioners would need to purchase to make their own prostheses would be prohibitive. Commercial KO sidebars are generally lighter and sleeker than the sidebars available to practitioners, and the joints are usually polycentric or anatomically correct. Knee brace companies generally vary the weight of their custom KO by changing the design and thickness of the thigh and calf sections; they use less weight for lighter versions. Some companies such as Becker Orthopedic (Troy, MI, U.S.A.), Otto Bock Health Care (Minneapolis, MN, U.S.A.), and United States Manufacturing Company (Poulsbo, WA, U.S.A.), distributors of sidebars and joints, offer a number of different styles and sizes to practitioners, but they are often considered too heavy or thick for incorporation into custom functional KO. These larger and heavier sidebars and clevis-style joints with reinforced plate steel at the joint mechanisms are most suitable for KO and knee/ankle/foot orthoses (KAFO) fabrication that aid chronic neuromuscularly challenged patients who require larger and heavier systems. Currently, steel joints with aluminum sidebars are the most common materials used in KO and KAFO. Metal sidebars and joints make up between one-third and two-thirds of the weight of typical knee orthoses.

Modern composite materials have been replacing metals in many applications, including orthotics and prosthetics. Composite materials are known to have good strength and stiffness properties, and because they are typically lighter than metals, their strength/weight and stiffness/weight ratios are particularly good. The fatigue resistance of composites is excellent when compared with aluminum. 1 A logical extension of the ubiquitous use of composite materials would be to embed the joint in a composite lay-up and eliminate sidebars in order to lighten the orthosis while retaining strength properties or stiffness properties or both. An orthotic fabrication platform, Composite Reinforcement Instead of Sidebars (CRIS), has been developed recently to make this possible. 2 The CRIS platform allows light and stiff orthoses without sidebars to be fabricated by practitioners in a potentially cost-saving method. The fabrication process lends itself to versatile designs, because the materials are laid-up, not contoured. There are no tooling costs for sidebars, and additional layers are applied according to the area of the patient that will be reinforced, the patient’s weight, activity level, and age. For instance, at an area of larger contours—just distal to the medial knee—more material can be applied to avoid bending failure. This area is often a fracture site in KAFO with aluminum sidebars, because the excessively contoured aluminum is weakened where it needs the most strength.

The potential for lightening and strengthening a KAFO is even greater than that for a KO because the ratio of sidebar weight to total orthosis weight is higher for a KAFO than for a KO. In past studies, patients with plastic/metal KAFO that were 15% lighter than all-metal KAFO were able to walk, on average, 17% faster at top speeds. 3 The increase in speed was even greater for small children, where the ratio of orthosis weight to patient weight was higher. 3 There have been some remarkable recent developments with automatic locking 4 and braking 5 for KAFO, but sidebars are still being used. Embedding the joint and removing sidebars would allow more room in the orthosis that would make it easier to add electronic components such as sensors, logic controls, servomotors, or power sources. The orthoses would also be more streamlined and less bulky without sidebars.

One of the main applications of KO is to provide support for patients with ACL injury, for whom the brace inhibits anterior tibial displacement (ATD). Many of these patients are active and would benefit from a lighter orthosis. The major objective of this work is to study the effectiveness of knee braces fabricated by the CRIS bracing platform in preventing ATD. It is hypothesized that knee orthoses produced using composite materials by the CRIS technique would be significantly lighter than commercial knee orthoses, and they could be even more effective than the commercial braces in limiting ATD. Mechanical tests of knee orthoses were conducted using a surrogate leg designed to reproduce ATD as it occurs in humans with ACL damage. Several CRIS orthoses were compared with orthoses from five knee-brace companies in the study.


All mechanical tests were performed on the surrogate leg that was constructed at Dartmouth College and modeled after one described by Liu et al. 6 The surrogate leg, shown in Figure 1, was mounted in a test fixture attached to the base of an Instron 8511 servohydraulic load frame test machine (Instron Corp., Canton, MA, U.S.A.). The rigid foam of the thigh and calf surrogate was covered with semi-rigid polyethylene foam and a custom ¼ inch neoprene sleeve that was 21 inches long (U.S. Orthotics, Tampa, FL, U.S.A.) to simulate muscle and soft-tissue surface hardness, following the recommendations of Liu et al. 7 The proximal end of the femoral component of the surrogate leg was fixed to the test machine at a knee flexion angle of 20 degres. The horizontal tibia was pivoted at the ankle and was free to translate at the knee joint, where it was guided by a low-friction linear bearing. Loads were applied to the midpoint of the tibia of the surrogate leg, forcing the tibial component of the knee to translate in the anterior direction. The magnitudes of force and displacement were measured continuously at the loading point using Instron’s data acquisition system (Instron MAX; Instron Corp.) until a maximum tibial translation (at the knee) of 25 mm (1 in) was attained. This is similar to the test methodology used by Liu et al. 6 to study the effectiveness of knee braces in controlling ATD.

Figure 1:
Diagram of the surrogate leg constructed for this test program. All dimensions are in mm.

On each of the testing days, initial tests were performed on the surrogate leg with no knee brace attached in order to determine the force-displacement characteristics of the surrogate leg. Tests were then done on each of the eight knee braces listed in Table 1, and several series of tests were done on each brace. Before a test series was conducted, a knee orthosis was put in position on the surrogate leg with the assistance of a certified Orthotist. The straps of the brace were tightened either to a force of 53.5 N (12 lbs), measured with a force scale, or to a force value deemed standard by the orthotist. The procedure for each test was to apply a continuously increasing load until a maximum displacement (tibial translation) of 25 mm (1 in) was attained; the load was then decreased to zero. This process was repeated five times during each test series, and at least three good test series were run on each brace. Therefore, there were at least fifteen loading tests performed on each knee brace (five repetitions in each of at least three test series). Data from each test included force and displacement at the point on the tibia where load was applied. The force was measured by the Instron’s load cell 2518-601 (Instron Corp.), displacement was measured using the Instron’s displacement transducer, which was integral to the Instron 8511’s load frame (Instron Corp.), and both sets of data were sampled continuously by the Instron data acquisition system (Instron MAX; Instron Corp.). The data files’ values were entered on an Excel (Microsoft, Redmond, WA, U.S.A.) spreadsheet, and the data were later analyzed using Excel (Microsoft). In the analysis, the forces and displacements were first converted to equivalent forces and displacements at the knee by using measured distances for the surrogate tibia to produce lever ratios. From the dimensions given in Figure 1, we found that the displacement at the knee is twice (610/305) the value of displacement measured by the Instron (Instron 8511; Instron Corp.); the equivalent force at the knee is half (305/610) the value measured by the Instron’s (Instron Corp.) load cell (F in Figure 1). The resulting forces and displacements at the knee during the test were then plotted. Tests of five braces (C, D, E, F, and G) were conducted on two different days to ensure repeatability of the test results.

Table 1:
Results of knee brace tests on Dartmouth's surrogate leg force (at knee) required to displace knee joint 25 mm (1 in)

Tests of the force-displacement characteristics of the surrogate leg without a knee orthosis attached were done before and after the knee brace tests. It was found that throughout the loading cycle there was a relatively consistent force of approximately 16 N (3.6 lb) required to translate the knee. During unloading the force dropped to approximately 13.1 N (2.95 lb). This force was most likely caused by friction in the knee-joint mechanism and ankle joint, and it was the same before and after each day’s knee brace tests.

Results of the first three load-unload cycles of a typical test series (for CRIS#1 knee orthosis A) are shown in Figure 2. It can be seen that the loading curve for the first test cycle is different from those of the following cycles, although the unloading curves are similar for all three cycles. It appeared that the straps of the brace were tightening during the first loading cycle, so the brace took up more of the applied force for that initial cycle than for later cycles, when the straps had settled into position. During those later cycles, the first portion of the force-displacement curve was flat, indicating that there was essentially no force required to displace the knee beyond the force inherent in moving the surrogate leg, so the knee brace was providing no resistance to motion. After the straps tightened, however, the brace started providing resistance to knee translation, which caused the force-displacement curve to bend upward. This same sequence of events was noted for all knee orthoses tested. That is, the force-displacement curves for the first loading cycle were different from those for later cycles, but the subsequent four load cycles had nearly identical force-displacement behavior. For that reason, the force-displacement curves for the last four loading cycles in each test series were considered representative of the behavior of the knee orthosis, and those curves were analyzed to determine the stiffness characteristics of the knee braces. The unloading curves for the knee braces were not analyzed.

Figure 2:
Force versus displacement for the first three loading cycles—test A-2.


Two pieces of information were determined from each loading curve: 1) the maximum force applied to the surrogate leg with each knee brace; and 2) the stiffness of the knee brace. The maximum force for a given knee orthosis was determined as the force at the knee required to produce 25 mm (1 in) of translation in the anterior direction at the knee.

Results for the different knee orthoses are summarized in Table 1. Analysis of those results showed the following:

  • Statistical analysis of the data (t-tests) showed that there was no significant difference (> 95% confidence) between the results with hand tightened straps and force-scale (53.5 N) tightened straps; results for brace E can be seen in Table 1. As a result, we treated the cases where either the force scale or the orthotist had tightened straps equally.
  • The CRIS knee orthoses could withstand the largest forces without excessive tibial translation, followed by the Townsend Original (Townsend Design, Bakersfield, CA, U.S.A.), the DonJoy Defiance #2 (dj Orthopedics, Inc., Vista, CA, U.S.A.), and CTi (Innovation Sports, Foothill Ranch, CA, U.S.A.).
  • The DonJoy Defiance #1 (dj Orthopedics, Inc.), Lenox Hill Avenger (Seattle Orthopedic Group Inc., Poulsbo, WA, U.S.A.), and GII Extreme (Generation II Orthotics, Richmond, British Columbia, Canada) orthoses were less effective in resisting tibial translation. The GENII (Generation II Orthotics) orthosis provided the least resistance to tibial translation.
  • B and H, two DonJoy braces (Defiance; dj Orthopedics, Inc.), were tested because the first such brace did not fit the surrogate leg properly. The better-fitting brace (H) provided more resistance to tibial translation, thus demonstrating the importance of proper fit of a knee orthosis.
  • When a strap was repositioned to the posterior from the anterior on the distal thigh section of CRIS orthosis A, the force required to displace the knee increased significantly. Orthosis D had both distal posterior and distal anterior thigh straps. All of the other braces had distal posterior thigh straps; brace E had one wide posterior thigh strap. The revised CRIS orthoses (A, with distal posterior strap, and G) were the most effective of all braces studied in resisting tibial translation.

The stiffness (force per unit displacement of the knee) was computed from the slope of the linear portion of the force-displacement curves for the last four loading cycles. The first custom CRIS knee orthoses tested was A. Figures 2 and 3 are using the tests done on A as an example of graphs done on all knee orthoses (A–H) in the experiment. There was a very good linear fit for the force versus displacement data (i.e., correlation coefficient r2 of 0.99 or greater for almost all cases). Therefore, all braces could be considered to be linearly elastic over the load and displacement range tested. The stiffness—resistance to tibial displacement when an anterior force is applied to the tibia—remained essentially constant both within a test and between tests for a given KO.

Figure 3:
Slope—best linear fit—of force-displacement curves for test A-2.

Stiffness measurement results for the eight knee braces are summarized in Table 2. It can be seen that the stiffness data correspond quite well with the maximum force data listed in Table 1. In particular, the two CRIS orthoses (A and G) proved to have the greatest stiffness, followed by orthosis F (Townsend Original; Townsend Design). Brace E (GII Extreme; Generation II Orthotics) was the least stiff. On a basis of stiffness per unit mass, orthoses A (CRIS #1) and G (CRIS #2) came out a factor of two better than the other braces, with brace G being the top-ranking owing to its lower mass (Table 2 and Figure 4). Orthosis E was the lowest, and the other braces had approximately equivalent stiffness/mass. The stiffness/mass results shown in Figure 4 indicate the benefits of producing sidebarless knee braces from light, stiff, composite materials.

Table 2:
Knee brace stiffness (anterior/posterior translation).
Figure 4:
Stiffness per unit mass of the knee braces tested. A = CRIS #1; B = DonJoy Defiance #1; C = CTi II; D = Lenox Hill Avenger; E = GII Extreme; F = Townsend Original; G = CRIS #2; and H = DonJoy Defiance #2.

In contrast to the maximum force results (Table 1), the stiffness results showed no significant difference for orthosis A between the tests done without a distal strap and those done with a distal strap. The reason is that stiffness is primarily a function of the brace design and materials, whereas the maximum force is also influenced by the displacement at which the straps begin to transfer force to the brace. Before the straps become taut against the leg, there is essentially no force taken up by the brace, and the tibia displaces without much required force. Then, when the straps become taut, force is transferred to the brace, and the force-displacement behavior is determined by the stiffness of the brace. This behavior is apparent in Figure 2 for orthosis A (with distal posterior strap). In our tests, orthosis A did not begin taking on significant force (defined here as a force of 3 lb = 13.4 N) until the knee had displaced an average of 14 mm with distal anterior strap (data not shown), but with a distal posterior strap the knee displaced an average of only 5.9 mm before 13.4 N of force was measured (Figure 2). It should be noted that with orthosis G, the knee displaced an average of 6.8 mm (data not shown) before 13.4 N was measured. The critical role played by the straps in the performance of knee braces is evident.

It may be of interest to compare the stiffness values measured in these tests with data obtained by Liu et al. 6 Four of the knee braces tested in our study were made by companies whose braces were also studied by Liu et al. It is not known whether the knee orthosis designs tested in that study were exactly the same as the ones we tested. Values of stiffness were determined using the data given in the Liu et al. paper, 6 and those values are given in Table 3, along with stiffness values from our study for the same manufacturer’s braces. It can be seen that our values are about a factor of 3 lower than those from the Liu et al. study. The reason for this may be primarily a different definition of force in that earlier study. Their paper is not sufficiently clear to enable the exact location of their force measurement to be determined, so their values cannot easily be changed to forces at the knee, as our values were. The relative ranking of the stiffness for the braces in the Liu et al. study are the same as ours. In both our tests, and in those of Liu et al., the braces made by Innovation Sports, Lenox Hill Brace Company and dj Orthopedics, Inc., had essentially the same stiffness, while the Townsend Design orthosis was measurably stiffer than the other three.

Table 3:
Comparison of stiffness data from these tests with stiffness data from Liu et al. 6


There are number of issues that need to be addressed about the relevance of this mechanical study to actual human experience. This was a successful proof of concept stage Phase I research study. Its total short-term purpose and focus is the construction of a lighter and stiffer KO using the prevention of anterior tibial displacement on a surrogate knee as measurement criteria. The long-term goal is to prelude the gathering of statistically relevant data in a Phase II human study for the purpose of scientifically proving that making a lighter and stiffer KO can offer benefits to patients.

Many professionals in the orthopedic field question the effectiveness of knee orthoses in preventing injuries for the ACL-involved patient. A common criticism is that the forces involved in sports that cause ACL injury are far greater than those that the KO can prevent. 8,9 Nonetheless, others have shown the benefits of such bracing. 6,7,10,11 A number of patients, which include orthotists and engineers, feel that the proprioceptive feedback offered by the KO acts as a cognitive and passive reminder that helps protect the knee from injury. 12 The knee is a complicated joint with the interaction of muscles, ligaments, cartilage, soft-tissue, and bones in a complex flexion/extension motion. These factors, along with vast individual differences in ACL patients, made it important to keep this early research with surrogates simple. The benefits of ACL bracing are hard to prove scientifically, but the market growth and success of the commercial brace companies indicate that patients and physicians must feel that there are benefits from wearing KO.

It is generally agreed that the ability of the KO to prevent anterior tibial translation is important for the success of an ACL KO. It is also generally accepted that a lighter orthosis is often preferential to patients. CRIS has succeeded at being a comparably light and stiff orthosis. It is likely that the fit of a KO is still the most important factor for patient acceptance, and this work showed that fit affects stiffness as well, but what a patient likes best and what performs best biomechanically are not always the same. For instance, a long orthosis made stiffer with more weight would prevent motion better than a lighter, shorter KO of the same material; however, the patient might choose the lighter and shorter KO for a number of reasons including cosmesis, weight, and comfort. 13 Ultimately, patient acceptance is the most important issue for long-term success for both patient and practitioner.

Another factor deemed to be an important feature by those in the orthopedic field when discussing KO is the ability of the mechanical joint to track the anatomical joint to prevent pistoning (parallel displacement of the orthosis), thereby enhancing fit. The mechanical joint used in the CRIS has a single axis (mono) at this time. Many believe that dual axis (poly) or anatomically correct joints are superior to mono. There is little scientific literature to support that hypothesis. The knee joint has very complicated kinematics that includes rotating, sliding, and gliding motion. Tests performed by Bahler did not find much difference between poly and mono joints, but poly did fare slightly better. 14 It was not stated which anatomically correct type joints were used, but they did not perform as well as mono or poly. The mechanics of pistoning in that study showed that mono, with a 16–22 mm backward displacement from the center of the femoral condyle axis, were better than the anatomically correct joints. 14 The best orthotic joints are poly with a backward displacement of 22 mm. Bahler states that the best functioning dual-axis knee joint would have the superior mechanical axis slightly anterior to the inferior because during flexion the anatomical knee rolls and glides inferiorly/posteriorly. He further states that experiments “show a center of rotation that remains practically unchanged up to almost 90 degrees; i.e., it scarcely rises or falls, it only moves backward.”14 In order to find out if anatomically correct joints really prevent pistoning, more research needs to be conducted.

It would be remiss to not point out that there are applications for which CRIS may not be preferred. A benefit of sidebars includes easy adjustability by a skilled technician or practitioner. The most common justifications for using metal systems are: 1) Severe fluctuating edema—often occurring in the lower limb; 2) Diabetic complications—neuropathy, insensate foot and poor skin integrity; and 3) Growth adjustments for children. In addition, aluminum sidebars can be contoured and recontoured when squaring the joint. Currently, with the CRIS wet lay-up method, adjustments can be made with a heat gun, but delamination can be a problem if not done carefully. With the CRIS, a square needs to be set on the positive mold before lamination to insure a free moving mechanical joint. Both aluminum and composite materials have their benefits and deficits.

The success that CRIS had in this test program perhaps relates to three major fabrication factors: 1) Continuous fiber fabrication method; 2) The placement of the angled carbon fiber sleeves; and 3) The embedding of the joint with fewer attachment sites compared to the sidebar KO. The materials are readily available and the procedure could be learned easily by those skilled in the art of composite resin laminations. Many improvements and innovations have been implemented since the initial development of CRIS, and it is anticipated that there will be more improvements in the continued development of CRIS.

Preimpregnated (prepreg) composite materials are thinner and stronger than the composites done in wet lay-up style. Currently, there is work being done at Rancho Los Amigos Hospital’s Research Department and at Otto Bock with prepreg in the fabrication of orthoses. The fabrication method of CRIS lends itself well to the use of prepreg. The cost of these materials is coming down as their use increases. The use of CRIS in a locking KAFO would offer potential benefits to the neuromuscularly challenged population. It is anticipated that the lighter weight and increased strength would allow these patients to walk faster while expending less energy. Also, CRIS prepreg KO could create a strong, light, and inexpensive off the shelf KO after positive molds are sized and prepared for prepreg layout. These are just a few of the research avenues that will be explored and tested using CRIS platform.

The feasibility of fabricating stiff and light KO has been accomplished, as it has been demonstrated in this study. Further mechanical tests are required in which medial/lateral forces and rotational torques are measured. Cyclic fatigue testing is also anticipated. A small preliminary human study at the Massachusetts General Hospital Biomotion Lab has been completed and the results are encouraging. The results of a subjective patient questionnaire as part of the Human Study were also positive. We still need to know whether patients will wear their CRIS or other orthoses or both and for how long during the day over an extended period they will be worn. This is the ultimate proof of acceptability


A mechanical test program was carried out in which the ability of knee braces to limit anterior tibial displacement in a surrogate model of the knee was determined. Knee orthoses from five different knee brace companies were compared with sidebar-less braces made from composite materials by a new orthotic fabrication methodology—CRIS. The orthoses made from composite materials were found to be the most effective of all braces studied in resisting tibial translation in the surrogate model. Their higher stiffness and lower weight gave them a considerably higher stiffness/weight ratio than the commercial knee braces with metal sidebars. The fit of the knee brace and the location of straps were also found to be important factors. As a result of their high resistance to surrogate anterior tibial displacement, low density, and high stiffness/weight ratio, knee orthoses made from modern composite materials could have considerable potential for patients with ACL damage.


The work reported here was supported by NIH Phase I SBIR #1R43HD36939.

T. Lunsford, CO, assisted in the design of the surrogate leg.


1. Lunsford TR. Strength of Materials in O&P Design. Alexandria, VA: American Academy of Orthotists and Prosthetists; 1996.
2. Glynn DW. US Patents #5,630,791 and #6,171,535.
3. Krebs DE, Edelstein JE, Fishman S. Comparison of Plastic/Metal and Leather/Metal Knee-Ankle-Foot Orthoses. Am J Phys Med Rehabil 1988; 67: 175–185.
4. Horton G. Field Facts: Horton Orthotic Lab ready to unveil new joint. O&P Business News 2001; 18: 46.
5. Kaufman KR. Energy-Efficient Knee Ankle Foot Orthoses: A Case Study. J Prosthet Orthot 1996; 8: 79–85.
6. Liu SH, Lunsford T, Gude S, Vangsness CT. Comparison of Functional Knee Braces for Control of Anterior Tibial Displacement. Clin Ortho 1994; 303: 203–10.
7. Liu SH, Daluiski A, Kabo JM. The Effects of Thigh Soft-Tissue Stiffness on the Control of Anterior Tibial Displacement by Functional Knee Orthoses. J Rehab Res Dev 1995; 32: 2–8.
8. Mortensen W, Forman K, Focht L, Daniel D. An In Vivo Study of Functional Orthoses in the ACL Disrupted Knee. Orthop Trans Res Soc 1988; 13: 520–8.
9. Beck C, Drez D Jr, Young J, Cannon W Jr, Stone M. Instrumental Testing of Functional Knee Braces. Am J Sports Med 1986; 14: 253–6.
10. Alexander H. In Search of the Perfect ACL Brace. U.S. Navy Clinical Investigation Program #90–48–2665, 1992:121.
11. Beynnon B, Pope M, Wertheimer C, et al. The Effect of Functional Knee Braces on the Anterior Cruciate Ligament In Vivo. J Bone Joint Surg Am 1992; 74: 1298–1312.
12. Wojtys E, Loubert P, Serafin S, Viviano D. Use of a Knee-Brace for Control of Tibial Translation and Rotation. J Bone Joint Surg Am 1990; 72: 1323–9.
13. Cawley P. Functional Knee Bracing for Skiing: A Review of Factors Affecting Brace Choice. Top Acute Trauma Rehabil 1988; 3: 73–81.
14. Bahler A. Fundamental Biomechanical Principles in the Orthotic Treatment of the Knee. J Prosthet Orthot 1992; 4: 157–63.

Knee orthosis; anterior tibial displacement; mechanical testing

© 2002 American Academy of Orthotists & Prosthetists