Development of an in vitro three dimensional loading-measurement system for long bone fixation under multiple loading conditions: a technical description
© Janicek et al; licensee BioMed Central Ltd. 2007
Received: 21 May 2007
Accepted: 24 November 2007
Published: 24 November 2007
The purpose of this investigation was to design and verify the capabilities of an in vitro loading-measurement system that mimics in vivo unconstrained three dimensional (3D) relative motion between long bone ends, applies uniform load components over the entire length of a test specimen, and measures 3D relative motion between test segment ends to directly determine test segment construct stiffness free of errors due to potting-fixture-test machine finite stiffness.
Intact equine cadaveric radius bones, which were subsequently osteotomized/ostectomized and instrumented with bone plates were subjected to non-destructive axial, torsion, and 4-point bending loads through fixtures designed to allow unconstrained components of non-load associated 3D relative motion between radius ends. 3D relative motion between ends of a 50 mm long test segment was measured by an infrared optical tracking system to directly determine its stiffness. Each specimen was then loaded to ultimate failure in either torsion or bending. Cortical bone cross-section diameters and published bone biomechanical properties were substituted into classical mechanics equations to predict the intact test segment theoretical stiffness for comparison and thus loading-measurement system verification.
Intact measured stiffness values were the same order of magnitude as theoretically predicted. The primary component of relative motion between ends of the test segment corresponded to that of the applied load with the other 3D components being evident and consistent in relative magnitude and direction for unconstrained loading of an unsymmetrical double plate oblique fracture configuration. Bone failure configurations were reproducible and consistent with theoretically predicted.
The 3D loading-measurement system designed: a) mimics unconstrained relative 3D motion between radius ends that occurs in clinical situations, b) applies uniform compression, torsion, and 4-point bending loads over the entire length of the test specimen, c) measures interfragmentary 3D relative motion between test segment ends to directly determine stiffness thus being void of potting-fixture-test machine stiffness error, and d) has the resolution to detect differences in the 3D motion and stiffness of intact as well osteotomized-instrumented and ostectomized-instrumented equine radii.
Biomechanical studies that appear in the literature typically use fixtures that constrain some components of relative 3D motion between ends of the test specimen, other than in the direction of loading, and/or do not apply a uniform component of load over the instrumented length of the specimen [1–8], and use test machine ram displacement to determine stiffness without correcting for potting-fixture-test machine (PFT) stiffness error [1, 2, 4, 6, 7, 9]. Direct measurement of relative 3D interfragmentary motions during in vitro testing has been performed [5, 7, 10–12], including those that have eliminated PFT stiffness [11, 12]. In vivo 3D interfragmentary motions have been measured using a transducer telemetry system , a motion capture system [6, 7, 9], and a 3D optical tracking system . To the authors' knowledge, combined use of 3D unconstrained fixtures with 3D optical tracking of test segment ends to eliminate PFT stiffness error has not appeared in human or veterinary literature.
The purpose of the investigation described in this paper was to design and verify the capabilities of an in vitro loading-measurement system that: 1) mimics in vivo unconstrained relative 3D motion between fracture-model (herein after referred to as "fracture") segments of long bone ends, 2) applies uniform compression, torsion, or bending moment loads over the entire length of the test specimen, 3) measures 3D relative motion between test segment ends to directly determine 3D stiffness components of intact and instrumented test segments, and 4) identifies the weakest aspects of an instrumented specimen during increased uniform loading over its entire length. The ultimate goal was to use this loading-measurement system to measure and compare the relative 3D biomechanical characteristics of intact and two Association for the Study of Internal Fixation (ASIF) techniques for repair of an oblique fracture involving the distal (inferior) diaphysis of an equine radius.
Loading fixture design, and potting
Proximal (superior) and distal steel box type loading fixtures were designed to compliment application of ASIF devices along the distal aspect of the lateral and cranial (anterior) radius surfaces with at least a 0.5 cm distance from potting material, and to have sufficient bone-potting interface to transmit loads to failure of the instrumented test segment.
Unconstrained 3D loading system design
An Instron 8821S (Instron, 100 Royall Street, Canton, MA) biaxial loading frame equipped with a 25 kN axial, ± 225 Nm torque load cell was used along with the loading fixtures described below to apply a 3D component of load uniformly over the length of the radius. Position control was used to move the global component at a constant rate of relative displacement between radius ends that corresponded to the applied load component. Assuming negligible friction in the lubricated joints, the fixtures were designed to produce no external constraints on the remaining (out of 6 total) 3D components of proximal relative to distal end displacement of the radius (herein after called "unconstrained"). The weight of the radius and fixtures were assumed to be negligible compared to the applied loads.
Axial compression was applied through hardened 25.4 mm diameter steel spheres centrally located in the end of each loading fixture base (Figure 2). Because of freedom of rotation of both ends about the sphere joints, there were no external constraints on proximal relative to distal end of radius: lateral-to-medial (LM) and CrCa translation (shear); and axial, LM and CrCa bending rotations.
3D relative inter-fragmentary motion measurement system
3D motion of the ends of the 50 mm long test segment were measured as load was applied by using a 0.01 mm resolution 3D infrared close focus Certus optical tracking system (NDI Optotrak, Waterloo, Ontario, Canada) to follow the light emitting diodes (LEDs) in the PVC rigid bodies attached at each end of the test segment using a #8–32 machine screw (Figure 2). A third rigid body was attached to a reference block on the base of the testing machine. The reference block had three miniature conical indentations located at known distances from the test machine's Z axis: one at the origin and one each on mutually perpendicular global X and Y axes. A 0.1 mm diameter tip portable digitizing probe was used to digitize the location of the three indentations on the reference block to define a global XYZ coordinate system, and 5 predetermined points around the fracture plane: one on each side of the lateral plate, one adjacent to cranial plate, and one each on the medial and caudal surface centerlines. Point digitization was done after each specimen was mounted in torsion fixtures with the medial side in the X direction. The digitized points and the optical tracking system's software was then used to define a local xyzd orthogonal coordinate axes associated with the rigid body attached at the distal end of the 50 mm long test segment, zd coincident with the radius' axis and xd in the medial direction. Similarly, a local xyzp coordinate axes was defined and associated with the rigid body attached at the proximal end. The local axes remained oriented with respect to and attached to the distal and proximal rigid bodies, thus tracking the 3D position of the radius' cross section at each end of the test segment as the specimen was placed in and loaded in each mode of loading.
Test segment stiffness determination and test system verification
To verify the capabilities of the 3D loading-measurement system, the stiffness associated with each mode of loading was measured for intact radius test segments (n = 20), and compared to the corresponding stiffness theoretically predicted with equations of classical mechanics. Each radius was then osteotomized, instrumented with ASIF techniques, and subjected to the same sequence of randomized 3-cycle non-destructive load tests (n = 20). Eight osteotomies were then converted to 7 mm ostectomies, re-instrumented, and subjected to the same sequence of 3-cycle non-destructive load tests. Representative results of the intact, osteotomized, and ostectomized tests for each load mode were compared to each other to determine: if the measurement system could resolve the primary component of 3D motion corresponding to direction of applied load from the other components, and if it had sufficient resolution to measure the stiffness of intact segments for comparison to that of fractured segments secured with different types of internal fixation. Following the last of these tests, each radius was subjected to either a 1-cycle torsion, LM bend or CrCa bend to failure test. The location and orientation of the fracture surfaces were observed and compared to the classical brittle fractures that would be produced by a pure component of applied load to help verify that the loading system is indeed applying the desired "pure" component of load.
Each intact radius was subjected to a 3-cycle non-destructive random sequence of load modes: axial compression, torsion, and CrCa and LM bending (n = 12), and in addition CaCr and ML bending (n = 8). Instron's RSBasLab software was used to apply load at a constant ram displacement rate (0.004 mm/sec axial compression, 0.25 deg/sec torsion, and 0.02 mm/sec bending), with load limits used to define ends of a cycle (100 to 4000 N axial compression, ± 50 Nm torsion, and 10 to 150 Nm bending). The 3D position of the rigid body LEDs were synchronously measured and stored along with the testing machine's ram position and applied load at sampling rates of 50 Hz axial, and 100 Hz torsion and bending. NDI Toolbench software was used to determine the 3D position of the proximal end of the test segment relative to the distal end for each test and to transfer this data along with applied load into an EXCEL file. EXCEL was used to plot the 3D components of relative position as a function of the component of applied load.
The component of relative displacement between test segment ends was plotted as a function of the corresponding applied load component. Measured stiffness was determined as the reciprocal of the slope of a straight line fit to this data between selected load ranges, typically having linear data, during the 3rd loading cycle: 200 to 4000 N for axial compression, 10 to 50 Nm for positive torque, -10 to -50 Nm for negative torque, 10 to 150 Nm for bending.
To compare the intact measured radius stiffness results (herein after called K3D) to those published by Hanson et al.  who used ram displacement in determining stiffness values (herein after called KH), the following was performed. Hanson et al's  axial and torsion tests were conducted on a 150 mm long test segment, thus the K3D results were multiplied by the ratio of test segment lengths (50 mm/150 mm) to obtain corresponding 150 mm segment length normalized values (herein after called K3DN) for direct comparison KH results. Hanson et al's  4-point bending tests were conducted with an apparatus having 11.5 mm between inner force application points and 101 mm between outer force application points, with both inner and outer forces applied directly to bone. Hanson et al.  did not report how relative angular displacement between cross sections was determined from the ram's axial displacement or between which pair of cross sections. Thus we converted our measured K3D values to those for a uniform bending segment of 11.5 mm and 101 mm by multiplying by the ratio of test segment lengths (50 mm/11.5 mm) = 4.35 and (50 mm/101 mm) = 0.50 respectively, to determined a range of K3DN values within which KH values should theoretically lie.
Theoretical segment stiffness
After testing to failure, each radius was transected 1 cm distal to the proximal rigid body allowing periosteal and endosteal cross section major and minor diameters to be measured with digital calipers. Two cross section shapes, hollow ellipse and hollow rectangle, were assumed to calculate a range of theoretically predicted stiffness values that would account for the non-uniform shape of the radius. The measured cross section diameters and published bone biomechanical properties were substituted into classical mechanics stiffness equations to predict its compression, torsion and bending stiffness values (Appendix).
3D loading-measurement system verification. Measured intact equine radii stiffness values for selected load ranges of the third loading cycle were observed to be the same order of magnitude as theoretically predicted values for compression, torsion, and 4-point bending.
Mean (± SD) stiffness of 50 mm long test segment
Axial Compression (N/mm)
+ Torsion (Nm/deg)
- Torsion (Nm/deg)
CrCa Bend (Nm/deg)
CaCr Bend (Nm/deg)
LM Bend (Nm/deg)
ML Bend (Nm/deg)
200 to 4000 N
10 to 50 Nm
-10 to -50 Nm
10 to 150 Nm
Measured using: 3D loading-measurement system
295000 (± 130977)
+430 (± 49)
-405 (± 57)
1007 (± 258)
885 (± 117)
1680 (± 395)
1452 (± 368)
Theoretical using: classical mechanics equations, cross-section measured dimensions, & two shapes
278000 (± 15261)
250 (± 26)
400 (± 49)
879 (± 94)
354000 (± 35277)
556 (± 57)
679 (± 83)
1490 (± 159)
Comparison of length normalized 3D loading-measurement system mean stiffness (K3DN) results to published conventional test machine load-measurement system stiffness (KH) results of intact equine radii. Differences in testing protocols are listed in the table. To be able to directly compare our 50 mm test segment stiffness (K3D) results, they were normalized to comparable values for test segment lengths used by Hanson et al.  by multiplying K3D by the ratio of test segment lengths: for axial and torsion (50 mm/150 mm), and for bending (50 mm/11.5 mm) and (50 mm/101 mm) to produce a range of K3DN values within which KH should lie.
Mean stiffness (K) of intact equine radius
Axial Compression (N/mm)
CrCa Bend (Nm/deg)
LM Bend (Nm/deg)
Measured using: 3D unconstrained fixtures
(3D optical) 34 cm test section 4-point bend, force applied on steel extensions to potted radius ends
3D optical tracking of the 50 mm test segment ends
K 3D (50 mm test segment)
K 3DN (150 mm axial & torsion segment)
(11.5 to 101 mm uniform bend segment)
4378 to 504
7305 to 840
Measured using: Constrained axial compression and torsion fixtures
(Hanson et al.) 1.15 cm test section 4-point bend (10.1 cm between outer force applicators), force applied directly on bone.
Test machine ram displacement
K H (150 mm axial & torsion segment)
(11.5 mm between inner, 101 mm between outer 4-point bend force application points)
Construct failure configurations were consistent with theoretical failure modes for brittle material (bone) and reproducible for both the torsion and bending load to failure tests. Torsion failure was consistently attributable to spiral fracture propagation that initiated perpendicular to the osteotomy or ostectomy at cranial screw holes. LM and CrCa bend failure initiated by cis-cortex fracturing through either the proximal or distal screw holes adjacent to the osteotomy or ostectomy, and progressed perpendicular to the longitudinal axis of the radius; with the exception of one ostectomy undergoing LM bend to failure experiencing substantial bone-plate interface separation due to lateral screw pull out of bone without any bone fracture. LM and CrCa bends produced visible osteotomy and ostectomy gap opening on the non-plated side due to it being in tension, and visible proximal and distal screw bending adjacent to the osteotomy or ostectomy. ML and CaCr bending produced visible ostectomy gap closure on the non-plated side due to it being in compression.
No failures occurred within the potted ends or of the fixtures themselves during torsional and bending loads to failure. No failures occurred through rigid body screw holes.
Measured stiffness (Ksm) of a long bone specimen is a function of parameters associated with morphologic/anatomical characteristics (material properties, cross section size and shape, length) which in turn are influenced by animal (age, sex, size, breed, diet, and training history), and parameters associated with testing protocol (specimen history/test conditions that effects its material properties [wet/dry, temperature, time from death to removal to storage, time in storage], test loading [strain] rate, relative stiffness of the potting-fixtures-testing machine [Kpft] to actual stiffness of the specimen [Ksa], fixture constraints on the relative 3D motion between specimen ends, and load range of data used for stiffness calculation). Hanson et al.  considered all but the later three when discussing why stiffness values from all modes of testing (compression, torsion, bending) intact equine tibia by McDuffee et al.  were 2.5 to 6 times greater, and contributed this to McDuffee's greater loading rates. McDuffee et al. [2, 15], Hanson et al. , and our compression testing was conducted in position control at constant ram velocity, with corresponding constant strain rate (ram velocity/exposed length of tibia or radius) = (0.58, 0.54), 0.00011, and 0.000018 (1/sec). McDuffee's compression strain rates were 527 and 491 times higher than Hanson's while ours' was 0.16 times lower. Similarly for torsion, (ram angular velocity/exposed length of tibia or radius) = (0.058, 0.054), 0.0017, and 0.00114 (deg/[mm*sec]); with McDuffee's being 34 and 32 times higher than Hanson's while ours' was 0.67 times lower. A quantitative comparison of strain rate for bending is not realistic due to load [2, 15] versus position control  used, and that applied bending moment (strain) rate is not uniform over the length of a specimen for 3-point bending, ranging from 0 at the outer supports to a maximum at the central load application point.
Panjabi et al.  reported that cortical bone longitudinal modulus of elasticity increases with increase in strain rate by a factor of 1.5 for strain rate of 1500 compared to 1.0 (1/sec), thus explaining in part the greater stiffness observed by McDuffee et al. [2, 15] compared to Hanson et al. . If strain rate were the predominant factor, our length normalized measured stiffness values should be lower than or similar to those reported by Hanson et al. , since our strain rates were lower.
The order of magnitude difference in our K3DN results compared to those reported by Hanson et al.  is likely due to their Kpft being relatively low compared to the Ksa of the specimen. When using test machine ram displacement to calculate Ksm of a specimen, ram displacement is the sum of that due to specimen deflection and PFT machine deflection, which includes bone crushing when applying bending loads directly to bone . The theoretical model is two springs in series , one with stiffness Ksa the other with Kpft, resulting in a measured stiffness Ksm = Ksa/(1 + Ksa/Kpft). Ideally Kpft is infinite (rigid PFT), thus Ksm = Ksa. The general rule-of-thumb when using ram displacement is to have Kpft ≥ 10*Ksa so that Ksa ≥ Ksm ≥ (1/1.1) Ksa and thus error is ≤ 9% when reporting Ksm for Ksa. Assuming K3DN = Ksa and KH = Ksm produces Ksa/Ksm ratios ranging from 7.5 to 19.6, which would require that Hanson's Kpft = 0.15 Ksa to 0.05 Ksa. This is low but possible; however, Cowin stated that with many testing machines, the stiffness of the bone specimen is greater than that of the load frame . Difference in Kpft may also have contributed to the higher stiffness reported by McDuffee et al. [2, 15] compared to Hanson et al. . Use of direct measurement of test segment end motion to determine its stiffness is not subject to PFT stiffness error, which is one reason we selected to use 3D optical tracking and is reason for using an extensometer when testing for material modulus of elasticity . The holes drilled in the radius to mount the LED rigid bodies were at the end cross sections of the test segment, thus having no effect on its stiffness.
Theoretically, in the absence of friction, the loading fixtures presented in this investigation allow 3D unconstrained components, of proximal relative to distal end motion of the radius, other than associated with the applied load component. Thus during loading, the ends of the specimen move into a relative 3D position that is dictated only by the 3D resistance of the (unsymmetrical or not) instrumented segment to the known (only) component of externally applied load such that mechanical equilibrium is achieved. The bending fixtures are the most likely to be subject to constraint by friction since the hardened cross bars were observed to indent the unhardened extension pipe surface during bending load to failure; the effect of this was reduced by periodically smoothing the pipe contact surfaces with a file.
Improvements could be made to reduce joint friction by using rollers incorporated into the 4-point bending load applicators . Use of a commercially available universal joint has been reported  and was considered but not used since its pivot center would be several centimeters above or below the proximal or distal end of the bone being tested, respectively. Since the applied axial load remains coincident with the axis through the center of the universal joints, this would create, as load was increased, an ever increasing non-physiological lateral displacement of the plate from the axis of loading and thus bending moment on the specimen. Our goal was to keep the center of the spheres as close as possible to the ends of the radius being tested, so that the line of action of the axial applied load would remain close to the central point on each end of the radius. The authors' feel that the loading fixtures used and described in this manuscript, particularly axial and torsion, are a significant improvement towards achieving clinically relevant 3D unconstrained "worst case" in vivo loading simulation, compared to the fixtures reported in the literature [1, 2, 4].
Another advantage of this loading fixture design was that the same potting of the ends could be used to test a specimen in all three modes of loading: axial compression, torsion and bending. In addition the bending fixtures were capable of applying uniform 4-point bending moment in any transverse plane over the entire length of the instrumented specimen with no point of load application applied directly to bone or instrumentation hardware, and is independent of specimen and fixture cross section size and material since the center pivot assures equal forces at the inner application points. Thus the weakest aspect of an instrumented segment can be located, in contrast to it being predisposed to be at the location of the highest applied bending moment [1, 2, 15]. Also it avoids the problem of applying loads to non-circular bone cross sections and failure due to bone crushing at the point of transverse force application .
Stiffness is typically determined as the slope of the "linear portion" of each load versus displacement curve [1, 2, 15] leaving the range of data used being subjectively determined and variable, which can have a notable effect on the numerical value obtained. To reduce the subjectivity, we used data over a consistent load range to determine stiffness for each loading modality. The load range was selected so that all tests would have a linear load versus displacement characteristic within the range. Our load versus test machine ram displacement curves typically had lower slope during the first loading cycle compared to 2nd and 3rd loading cycles (the later two being similar), due to settling of bone-potting-fixture interconnections during the first cycle . McDuffee et al. [2, 15] and Hanson et al.  obtained stiffness from single cycle to failure tests, which could be a factor in the variability of and lower stiffness values observed.
The theoretical stiffness equations are based on the assumption that cross sections remain plane as load is applied, and thus corresponding experimental stiffness determination depends upon accurate 3D measurement of the position of these planes. For axial and torsional testing the initial position of the LEDs was pointing directly towards the 3 camera optical measuring head (Figures 2, 3) but were oriented at approximately 45 degrees for bending (Figures 4, 5). Accuracy in detecting the location of individual LEDs is known to deteriorate as their angle deviates from pointing directly at the measuring head . This may explain why the average measured axial and torsional stiffness values were within the predicted range, while the average measured bending stiffness values were 0.97 to 1.5 times greater than the largest theoretically predicted stiffness for the four bending modes. This also might be an explanation for the occasional saw tooth type characteristic in some bending moment versus relative angular displacement graphs to which we used regression to fit a straight line to get the general trend and thus stiffness. Another possibility is that the single point of attachment to the radius of each rigid body did not produce an accurate representation of the 3D motion of the cross section plane. A recommended option would be to attach two rigid bodies to a C-ring with three pointed machine screws to make cortical bone contact at three points on the cross section plane. This would also eliminate the LED rigid body mounting hole's stress riser and potential for fracture initiation at this site, which did not occur in our tests to failure.
Long bone is known to be anisotropic and heterogeneous (differing in diaphyseal and epiphyseal properties), and is primarily elastic at low deformation rates [16, 17]. The principal stress is longitudinal for compression and bending load modes. Published low strain rate longitudinal modulus of elasticity E determined by machine testing ranged from 17,000 to 22,600 (N/mm2) for human and bovine femur and tibia . Torsional stiffness is a function of the longitudinal-circumferential shear modulus G, with published machine testing determined values ranging from 3300 to 5000 (N/mm2) for human and bovine femur and tibia . Thus use of E = 18,000 and G = 4,615 for the short 50 mm long intact diaphyseal test segment produced representative theoretical stiffness values for use as a magnitude accuracy comparison reference.
In conclusion, assuming negligible friction, the 3D loading-measurement system described in this manuscript: a) mimics unconstrained relative 3D motion between radius ends that occurs in clinical situations, b) applies uniform compression, torsion, and 4-point bending loads over the entire length of the test specimen, c) measures interfragmentary 3D relative motion between test segment ends to directly determine stiffness thus being void of PFT machine stiffness error, and d) has the resolution to detect differences in the 3D motion and stiffness of intact as well osteotomized-instrumented and ostectomized-instrumented equine radii. It is the authors' opinion that the 3D loading-measuring system described in this manuscript is capable of creating "worst case" clinically relevant loading conditions, and accordingly has the capacity to detect location and type of instrumented segment modes of failure and has the ability to produce more accurate stiffness results that are independent of PFT machine stiffness error.
Axial compression stiffness: KC = AE/L(N/mm) (1)
Torsional stiffness: KT = (1.745 * 10-5)*GK/L (Nm/deg) (2)
LM & ML bending stiffness: KLM = (1.745 * 10-5)*EIyy/L (Nm/deg) (3)
CrCa & CaCr bending stiffness: KCC = (1.745 * 10-5)*EIxx/L (Nm/deg) (4)
Dp, De = Major diameter periosteal, endosteal respectively (mm)
dp, de = Minor diameter periosteal, endosteal respectively (mm)
A = Cross section area (mm2)
= (π/4) * (Dpdp - Dede) for hollow ellipse (5a)
= (Dpdp - Dede) for hollow rectangle (5b)
Iyy = Area moment about y axis (mm4)
= (π/64) * (Dp 3dp - De 3de) for hollow ellipse (6a)
= (1/12) * (Dp 3dp - De 3de) for hollow rectangle (6b)
Ixx = Area moment about x axis (mm4)
= (π/64) * (Dpdp 3 - Dede 3) for hollow ellipse (7a)
= (1/12) * (Dpdp 3 - Dede 3) for hollow rectangle (7b)
K = Polar moment type term (mm4)
= (π/16)* [(Dp 3dp)(1-q4)]/[Dp 2 + dp 2] for hollow ellipse (8a)
= (1/12)* [(Dpdp*(Dp 2+dp 2)) - (Dede*(De 2+de 2))] for hollow rectangle (8b)
q = ratio of major endosteal/periosteal diameters = (De/Dp)
L = Length of the test segment = 50 mm
E = Cortical bone modulus of elasticity = 18,000 (N/mm2)
Range in literature (8,900 ≤ E ≤ 42,000)
G = Cortical bone shear modulus = 4,615 (N/mm2)
Range in literature (3,500 ≤ G ≤ 8,710)
Association for the Study of Internal Fixation
Our intact measured stiffness results
Our intact 50 mm long segment stiffness results normalized to Hanson et al test specimen length
Hanson et al's ram displacement stiffness values
Potting-fixture-test machine stiffness
Actual stiffness of test specimen
Measured stiffness of test specimen
Light emitting diode
Local orthogonal axes associated with proximal test segment end
Local orthogonal axes associated with distal test segment end
Supported by the University of Missouri Comparative Orthopaedic Laboratory with financial assistance provided from the E. Paige Laurie Equine Endowed Program in Lameness and the University of Missouri Department of Orthopaedic Surgery to purchase the 3D optical tracking system. Presented in abstract form at the 2nd World Veterinary Orthopedic Congress/33rd Annual Conference of the Veterinary Orthopedic Society, Keystone, CO, 2006.
- Hanson PD, Markel MD, Vanderby R: Diaphyseal structural properties of equine long bones. Am J Vet Res. 1995, 56: 233-240.PubMedGoogle Scholar
- McDuffee LA, Stover SM, Taylor KT: An in vitro biomechanical investigation of the mechanical properties of dynamic compression plated osteotomized adult equine tibiae. Vet Surg. 1997, 26: 126-136.View ArticlePubMedGoogle Scholar
- Hopper SA, Schneider RK, Ratzlaff MH, White KK, Johnson CH: Effect of pin hole size and number on in vitro bone strength in the equine radius loaded in torsion. Am J Vet Res. 1998, 59: 201-203.PubMedGoogle Scholar
- Radcliffe RM, Lopez MJ, Turner TA, Watkins JP, Radcliffe CH, Markel MD: An in vitro biomechanical comparison of interlocking nail constructs and double plating for fixation of diaphyseal femur fractures in immature horses. Vet Surg. 2001, 30: 179-190. 10.1053/jvet.2001.21396.View ArticlePubMedGoogle Scholar
- Bernardé A, Diop A, Maurel N, Viguier E: An in vitro biomechanical comparison between bone plate and interlocking nail. Vet Comp Orthop Trauma. 2002, 15: 57-66.Google Scholar
- Klein P, Schell H, Streitparth F, Heller M, Kassi JP, Kandziora F, Bragulla H, Haas NP, Duda GN: The initial phase of fracture healing is specifically sensitive to mechanical conditions. J Orthop Res. 2003, 21: 662-669. 10.1016/S0736-0266(02)00259-0.View ArticlePubMedGoogle Scholar
- Kaspar K, Schell H, Seebeck P, Thompson M, Schütz NP, Duda GN: Angle stable locking reduces interfragmentary movements and promotes healing after unreamed nailing. Study of a displaced osteotomy model in sheep tibia. J Bone Joint Surg Am. 2005, 87: 2028-2037. 10.2106/JBJS.D.02268.View ArticlePubMedGoogle Scholar
- Higgins TF, Pittman G, Hines J, Bachus KN: Biomechanical analysis of distal femur fracture fixation: fixed-angle screw-plate construct versus condylar blade plate. J Orthop Trauma. 2007, 21: 43-46. 10.1097/BOT.0b013e31802bb372.View ArticlePubMedGoogle Scholar
- Schell H, Epari DR, Kassi JP, Bragulla H, Bail HJ, Duda GN: The course of bone healing is influenced by the initial shear fixation stability. J Orthop Res. 2005, 23: 1022-1028. 10.1016/j.orthres.2005.03.005.View ArticlePubMedGoogle Scholar
- Gardner TN, Evans M, Kyberd PJ: An instrumented special linkage for monitoring relative three-dimensional motion between fracture fragments. J Biomech Eng. 1996, 118: 586-593.View ArticlePubMedGoogle Scholar
- Duda GN, Kirchner H, Wilke HJ, Claes L: A method to determine the 3-D stiffness of fracture fixation devices and its application to predict inter-fragmentary movement. J Biomech. 1998, 31: 247-252. 10.1016/S0021-9290(97)00115-2.View ArticlePubMedGoogle Scholar
- Kassi JP, Hoffmann JE, Heller M, Raschke M, Duda GN: Evaluating the stability of fracture fixation systems: mechanical device for evaluation of 3-D stiffness in vitro. Biomed Tech. 2001, 46: 247-252.View ArticleGoogle Scholar
- Claes L, Augat P, Suger G, Wilke HJ: Influence of size and stability of the osteotomy gap on the success of fracture healing. J Orthop Res. 1997, 15: 577-584. 10.1002/jor.1100150414.View ArticlePubMedGoogle Scholar
- Maletsky LP, Sun J, Morton NA: Accuracy of an optical active-marker system to track the relative motion of rigid bodies. J Biomech. 2006, 40: 682-685. 10.1016/j.jbiomech.2006.01.017.View ArticlePubMedGoogle Scholar
- McDuffee LA, Stover SM, Taylor KT, Les CM: An in vitro biomechanical investigation of an interlocking nail for fixation of diaphyseal tibial fractures in adult horses. Vet Surg. 1994, 23: 219-230.View ArticlePubMedGoogle Scholar
- Panjabi MM, White AA: Material properties. Biomechanics in the Musculoskeletal System. Edited by: Panjabi MM, White AA. 2001, New York: Churchill Livingstone, 167-190.Google Scholar
- Cowin SC: Bone Mechanics. 1989, Boca Raton: CRC PressGoogle Scholar
- Florin M, Arzdorf M, Linke B, Auer JA: Assessment of stiffness and strength of 4 different implants available for equine fracture treatment: A study on a 20° oblique long-bone fracture model using a bone substitute. Vet Surg. 2005, 34: 231-238. 10.1111/j.1532.950X.2005.00035.x.View ArticlePubMedGoogle Scholar
- NDI Optotrak Certus Users Guide, revision 1.0. 2003, 63-
This article is published under license to BioMed Central Ltd. This is an Open Access article distributed under the terms of the Creative Commons Attribution License (http://creativecommons.org/licenses/by/2.0), which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.