Biomechanical Comparison of Three Locking Compression Plate Constructs from Three Manufacturers under Cyclic Torsional Loading in a Fracture Gap Model

Abstract

Objective The aim of the study was to compare the stiffness and cyclic fatigue of locking compression plate constructs from three manufacturers, DePuy Synthes (DPS), Knight Benedikt (KB), and Provet Veterinary Instrumentation (Vi), under cyclic torsion.

Methods The constructs of DPS, KB, and Vi were assembled by fixing a 10-hole 3.5-mm stainless steel locking compression plate 1 mm away from a validated bone model with a fracture gap of 47 mm. The corresponding drill guides and locking screws were used. Three groups of six constructs were tested in cyclic torsion until failure.

Results There was no significant difference in initial stiffness between DPS constructs (28.83 ± 0.84 N·m/rad) and KB constructs (28.38 ± 0.81 N·m/rad), and between KB constructs and Vi constructs (27.48 ± 0.37 N·m/rad), but the DPS constructs were significantly stiffer than the Vi constructs. The DPS constructs sustained the significantly highest number of cycles (24,833 ± 2,317 cycles) compared with KB constructs (16,167 ± 1,472 cycles) and Vi constructs (19,833 ± 4,792 cycles), but the difference between KB and Vi constructs was not significant. All constructs failed by screw damage at the shaft between the plate and the bone model.

Conclusion DPS constructs showed superior initial torsional stiffness and cyclic fatigue life than Vi constructs, whereas KB and Vi constructs shared comparable results. Further investigation is required to assess the clinical significance of these biomechanical differences.

Keywords

locking plate - cyclic loading - torsion - fatigue life - biomechanical performance of implants

Introduction

Bridging plate osteosynthesis has been widely used for treating long bone fractures in humans and small animals. Locking compression plates (LCP) are particularly useful and versatile for bridging osteosynthesis because they provide angular stability, require minimal contouring, and allow the flexibility of using a combination of standard cortical screws and locking screws.[1] [2] [3] They are also advocated as biological internal fixations with the design of undercut surface so that the soft-tissue environment and bone vascularity can be preserved.[4]

As locking plates grow in popularity among veterinary practices, there are increasing numbers of manufacturers of LCP systems with different designs of the plates and screws, which can potentially affect their mechanical properties. Three of the common manufacturers in the authors' region are DePuy Synthes (DPS; New South Wales, Australia), Knight Benedikt (KB; New South Wales, Australia) and Provet Veterinary Instrumentation (Vi; New South Wales, Australia). To the authors' knowledge, there is no literature available on comparison between LCP systems from these manufacturers, and the choice of implant manufacturers are mainly based on cost and surgeon preference.

Mechanical testing of orthopaedic implants commonly isolates forces that a fracture may experience. These forces include bending, torsion, and axial compression. Distribution of these forces on bone–plate constructs varies in clinical situation, and it depends on the anatomy of the bone, the phase of the gait cycles, and the size and activity style of the veterinary patients.[5] [6] [7] [8] [9] [10] [11] Strain distribution on animal models has been investigated by previous studies,[8] [9] [11] which show that bending and torsional forces are both major contributing forces on long bones under physiological load compared with axial compression. However, the complexity of these forces in clinical situation makes definitive replication of these forces in biomechanical testing impossible. The lack of standardized testing protocols also makes direct comparisons of results between studies challenging. These factors then limit extrapolation of biomechanical test results to clinical settings. Nevertheless, it remains important that the isolated forces on implant constructs are investigated so that problem areas of fixation can be recognized in relation to the weakness of constructs under specific circumstances. The majority of the previous studies comparing locking plate systems focus on testing of the plates and their failure. However, increasing incidence of screw failure in LCP systems was seen by the authors, and previous biomechanical studies reported relatively more screw failure scenarios when LCP constructs were tested under torsion.[12] [13] [14] [15] The purpose of this study was to compare the biomechanical behavior of LCP constructs of three manufacturers, including DPS, KB, and Vi, during fatigue testing in torsion until failure. We hypothesized that the constructs of DPS, KB, and Vi would have no significant difference in torsional stiffness and cycles to failure.

Materials and Methods

Bone Model

A fourth-generation short fiber filled epoxy hollow cylinder bone model (SawBones, Vashon, Washington, United States), 83 mm in length, 3-mm wall thickness, and 20-mm outer diameter, was used to stimulate bone segments.

Plates and Screws

LCP constructs from DPS, KB, and Vi were tested. The plates were 10-hole 3.5-mm stainless steel LCP and the screws were 3.5-mm self-tapping locking screws, 28 mm in length. Each plate was secured with its corresponding screws from the same manufacturer.

Construct Assembly

Two bone model segments were aligned and secured in a custom platform. A gap of 47 mm was made between the two segments to simulate a clinical comminuted fracture where LCP constructs were used in a bridging fashion. This gap was designed to allow the plate to center over the fracture gap with equal numbers of screw holes on each side. A metal spacer of 1-mm thickness was used to separate the plate from the bone surrogate. The plate was then centered and secured over the spacer-cylinder unit using clamps on each side. A corresponding locking drill guide was used for drilling. Three screws were placed in the outermost plate holes on each segment and they were hand-tightened to 1.5 N·m using the torque limiter supplied by the corresponding manufacturers. Correct insertion and appropriate seating of the locking screws were checked visually. A screw was considered as being incompletely seated when at least one thread of the screw head was visible above the surface of the plate. If the position of the screw was considered inappropriate despite multiple attempts, the construct would be discarded. No implant was reused for construct assembly or mechanical testing. Each end of the construct was potted in polymethyl methacrylate and polyvinyl chloride pipe. Two crossed Kirschner wires were placed through each end of the construct to ensure there was no delamination between the layers during testing. Six constructs per group were tested.

Mechanical Testing

Cyclic torsion tests on the constructs were performed using an electric linear fatigue testing machine (ElectroPuls E1000, Instron, Norwood, Massachusetts, United States) with a dynamic load cell of ± 1kN and ± 25 N·m. Each end of the construct was centered and rigidly fixed to the machine using custom-designed concentric clamps so that the construct was oriented vertically and the axial rotation was made through the center of the shaft of the bone surrogate ([Fig. 1]). Torque and axial load on the construct were made zero before testing, and axial load during torsion was controlled among constructs. Constructs were tested under cyclic torsion in displacement control with a sinusoidal waveform inducing an angle of 0 to +0.218 rad at a rate of 4 Hz until catastrophic failure. After three quasi-static load–unload tests to condition the construct, a quasi-static load–unload test was performed with an angle of 0 to +0.218 rad at a speed of 0.02 rad/s and it was repeated for 1,000 cycles followed by an acquisition cycle using the same parameters. The conditioning cycles, loading test cycles, and the acquisition cycle were repeated in 1,000-cycle blocks at a rate of 4 Hz until catastrophic failure.

Fig. 1 (A) Photograph of the mechanical setup designed for cyclic torsion, with a dynamic load cell of ± 1 kN and ± 25 Nm. A pair of custom-made concentric clamps (B) was attached to each end of the testing construct (C) so that axial rotation was made through the center of the sawbone shaft. The lower clamp was fixed to the INSTRON base (D) via an XY translation stage to remove any extra loading.

Data Acquisition

The performance of each construct was described by its initial torsional stiffness, number of cycles to failure, and failure modes. For each quasi-static load–unload test, the angle of rotation, torque, and displacement data were recorded throughout the test at a rate of 1,000 Hz using a dynamic data acquisition software (WaveMatrix2, London, UK). Torque versus angular displacement curves were generated from each set of quasi-static test data using commercially available spreadsheet software (Microsoft Excel, Sydney, Australia). Torsional stiffness was calculated from the slope of the linear portion of the torque–angular rotation loading curve, and it was expressed as Newton meters per radian (N·m/rad). The initial torsional stiffness of a construct was taken from the final conditioning cycle at the beginning of the test. Failure initiation was determined by the first significant drop in stiffness visually and mathematically using the formula from the study of Bilmont and colleagues[12]:

where D is the difference between two consecutive slopes of the stiffness–cycle curve, S_n is the stiffness for the cycle n (C_n ), and S_{n + 1} is the stiffness for the cycle n + 1 (C_{n + 1}). This difference was plotted against the number of cycles. Catastrophic failure of a construct was defined as any visual damage or deformation of the construct, or any visual abrupt drop in stiffness on the curve.

Statistical Analysis

The initial torsional stiffness and number of cycles to failure initiation and complete failure of the constructs were reported as mean ± standard deviation (SD). Statistical analysis was performed using a one-way analysis of variance (ANOVA) with post hoc Tukey's honestly significant difference test for comparison of stiffness and number of cycles to failure initiation and complete failure among the three manufacturer groups. Statistical significance was set at p < 0.05.

Results

The locking screws of DPS, KB, and Vi were all made of medical grade 316LVM stainless steel with an outer diameter of 3.5 mm, a core diameter of 2.9 mm, and a conical double threaded screw head. The specifications of the LCP are summarized in [Table 1]. Since specific details of the implants were considered proprietary by the manufacturers, some dimensions of the implants were measured by the authors. The design and shape of the combi-holes and undercuts of the plates were noticeably different ([Fig. 2]). The undercut of the Vi plate overlapped more with the locking screw hole compared with those of the DPS and KB plates ([Fig. 3]). Although all screw heads from the three manufacturers sat within the plate, there were less threads within the screw hole of the Vi plate compared with those of the DPS and KB plates ([Fig. 4]).

Table 1 Specifications of locking compression plates of DPS, KB, and Vi

	Length (mm)	Width (mm)	Height (mm)	Distance between centers of locking holes (mm)	Material	Maximum length of undercut (mm)[a]	Maximum height of undercut (mm)[a]	Length of combi-hole (mm)[a]
DPS	131	11	3.3	13	316LVM stainless steel	6	1.5	7.5
KB	137	11	3.3	13	316LVM stainless steel	6	1.5	7.5
Vi	133	11	3.3	13	316LVM stainless steel	8.5	1.5	8.5

Abbreviations: DPS, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation (Vi).

^a Data measured by the authors.

Fig. 2 Photograph of the undersides of the locking compression plates of DPS, KB, and Vi. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

Fig. 3 Photograph of the positions of the screws relative to the undercuts of the plates of DPS, KB, and Vi. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

Fig. 4 Photograph of the combi-holes of the locking compression plates of DePuy Synthes, Knight Benedikt, and Provet Veterinary Instrumentation, showing the difference in number of threads within the locking screw hole. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

There was not a significant difference in the initial torsional stiffness between the DPS (28.83 ± 0.84 N·m/rad) and KB (28.38 ± 0.81 N·m/rad) constructs (p = 0.53), or between the KB and Vi (27.48 ± 0.37 N·m/rad) constructs (p = 0.10; [Fig. 5]). However, the DPS constructs were significantly stiffer than the Vi constructs (p = 0.01).

Fig. 5 Histogram of the mean stiffness for the DePuy Synthes, Knight Benedikt, and Provet Veterinary Instrumentation constructs loaded under cyclic torsion. Error bars represent standard deviation. Constructs that share the same lowercase letter are significantly different. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

Constructs from all three manufacturers exhibited biphasic evolution of stiffness over time ([Fig. 6]). Stiffness remained stable until failure initiation, and a lower stiffness was maintained until complete failure, which was characterized by an abrupt drop in stiffness. Failure initiation occurred at the latest in the DPS constructs (11,833 ± 2,927 cycles; p = 0.005), while failure initiation did not differ significantly between the KB (7,833 ± 983 cycles) and Vi (7,833 ± 753 cycles) constructs (p = 0.9).

Fig. 6 Illustration of the evolution of stiffness over time for the constructs of DePuy Synthes (diamond), Knight Benedikt (square), and Provet Veterinary Instrumentation (triangle). The arrows indicate failure initiation and the stars show complete failure. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

The number of cycles to failure was the highest in the DPS constructs (24,833 ± 2,317 cycles), followed by the Vi constructs (19,833 ± 4,792 cycles) and KB constructs (16,167 ± 1,472 cycles; [Fig. 7]). The DPS constructs sustained a significantly higher number of cycles to failure compared with both the KB (p = 0.001) and Vi groups (p = 0.03), but there was no significant difference between the KB and Vi groups (p = 0.15).

Fig. 7 Histogram of the mean cycles to failure for the DePuy Synthes, Knight Benedikt, and Provet Veterinary Instrumentation constructs loaded under cyclic torsion. Error bars represent standard deviation. Constructs that share the same lowercase letter indicate no significant difference. Synthes, DePuy Synthes; KB, Knight Benedikt; Vi, Provet Veterinary Instrumentation.

All constructs failed by screw breakage at the free portion of the shaft between the plate and the bone model at catastrophic failure ([Fig. 8]). There was no failure of the bone surrogate or deformation of the plate noticed grossly at the end of testing. None of the implants were discarded due to improper construct assembly.

Fig. 8 Photograph demonstrating the mode of failure. Fracture of screws occurred at the free portion of the shaft between the plate and the bone surrogate.

Discussion

The results of our study showed that the DPS constructs withstood more cycles than the KB and Vi constructs. Although the DPS constructs had the highest initial torsional stiffness, there was no significant difference in initial torsional stiffness between the DPS and KB constructs. The torsional stiffness and cycles to failure of the KB and Vi constructs were not significantly different.

The construct stiffness presented in this study represents the initial fixation characteristic of the plate–bone constructs. With loading under torsion along the axis of the bone surrogate, a major bending moment was created at the free portion of the screw shaft, especially with the 1-mm offset of the plate from the bone; while the plate was likely subject to torsional moment. The overall construct stiffness is likely determined by the combination of bending stiffness of the screw and the torsional stiffness of the plate. On the other hand, since none of the plate showed any elastic deformation and all the constructs consistently failed by screw breakage at the free portion of the screw shaft between the plate and the bone surrogate, this suggested that the main determining factor for construct failure in this study was likely the fatigue life of the screw. The free unsupported part of the screw shaft was subject to high stress and greater deformation due to the lever arm effect under loading, thus leading to screw breakage. Given the biphasic evolution of stiffness of the constructs, the mechanism of screw failure in this study was consistent with previous studies[16] [17]: the construct fails under cyclic loading as a result of defect accumulation, crack initiation, and propagation of crack when the number of cycles increases.

Previous studies have demonstrated that multiple factors affect the stability and fatigue life of locking constructs including plate–screw density, plate span ratio, working lengths of plate and screw, insertion angle of locking screws, and plate–bone distance.[2] [3] [12] [14] [18] [19] [20] [21] [22] These factors were kept constant in this study, and the testing setup was designed to represent the clinical scenario. First, the number and placement of the screws on the plates were kept constant among the manufacturer groups, thus making the plate–screw densities of the constructs the same. Although no guideline is currently available in veterinary surgery regarding the optimal number of locking screws per fragments for comminuted fractures, the guidelines for clinical application of LCP in human literature[2] demonstrate that a plate span ratio greater than 2 to 3, and a plate–screw density of 0.4 to 0.5 are recommended for comminuted fracture repair. Previous studies[12] [23] [24] also show that constructs with three bicortical locking screws per fragment have a significantly higher stiffness and fatigue life compared with those with two screws per fragment. Therefore, in this study, a setup of three locking screws per fragment with a plate–screw density of 0.6 was chosen in this simulated comminuted fracture because this configuration is likely to be preferable clinically. Second, increasing the working length of the plate has been demonstrated to have a negative impact on construct rigidity and survival for large or comminuted fracture gaps.[2] [3] [20] [25] Since the plates were centered over the fracture gap with equal numbers of empty screw holes on either side and all three manufacturers' plates have the same distance between locking screw holes, working lengths of the plates of the three groups were the same despite a slight difference in plate length. Third, working lengths of the screw were kept constant among all the constructs. The shafts of the locking screws of the three manufacturers also share the same dimensions, thus making the area moment of inertia of the screws the same. Finally, a plate–bone distance of 1 mm was chosen in this study because the LCP are not always applied flush to the bone in clinical situations and this distance has been demonstrated to have an insignificant impact on construct stability.[14]

Despite the aforementioned factors of the construct being kept constant among three groups, significant differences in results were still present in this study. The cause of these differences in results is unknown, but it could be associated with the material and design of the implants as suggested by previous studies.[19] [26] [27] [28] [29] First, the difference in the design and shape of the LCP ([Fig. 2]) could have affected the geometric shape and cross-sectional area of the plates. Consequently, differences in the polar moment of inertia and area moment of inertia of the composite arose, resulting in differences in torsional and bending stiffness of the constructs, respectively. However, since the LCP from the three manufacturers share the same maximum height of the undercut, overall width, and height, the area and polar moments of inertia of the three LCP should be relatively similar. Unfortunately, since detailed dimensions of the implants are considered proprietary by the manufacturers, the required information for the exact geometric calculation was not available at the time of this study. The significance of the design of more overlapping undercut over the screw hole on the Vi plate ([Fig. 3]) was unknown. However, the authors speculated that this might reduce the cross-sectional area at that level, causing a reduction in the area moment of inertia and polar moment of inertia. This can lead to a decrease in bending and torsional stiffness, especially if stress is applied at that level. In addition, the lesser numbers of threads in the screw hole of the Vi plates ([Fig. 4]) can potentially increase the risk of off-axis insertion of locking screw and screw head loosening due to the reduced thread engagement. Previous studies[18] [22] demonstrate that off-axis insertion of locking screws with cross-threading causes a significantly compromised mechanical stability to the locking plate–screw interface by reducing failure load by 50% when deviation in angle of insertion is more than 5 degrees. Since screw head unlocking was not noticed in this study, the authors believe that the difference in the screw hole design had a minimal impact on the results of this study. However, it is recommended to insert locking screws perpendicular to the plate using a drill guide and to contour the plate accordingly in clinical scenarios. Finally, although all the screws from the manufacturers have the same area moment of inertia and all the implants used in this study are reported to be made of medical grade 316LVM stainless steel, there could be a difference in the processing and cold working of the stainless steel, which could influence the mechanical properties of the implants.[30] Assessment of the quality of the stainless steel was out of the scope of this study and the associated information was unavailable from the manufacturers.

The DPS constructs in this study showed comparable results to those in the study of Bilmont and colleagues,[12] which had similar mechanical testing setup with a lower rate of cyclic torsion. The DPS constructs in both studies demonstrated a biphasic evolution of stiffness with comparable failure initiation and cycle to failure. In Bilmont and colleagues,[12] approximately 40% of the constructs failed with screw head unlocking, while no failure was observed in this study. This could be explained by the fact that all screws in the current study were adequately seated before testing, while 9% of the screws in the study by Bilmont and colleagues[12] were not completely seated. Screw head unlocking due to inadequately seated locking screws has been previously reported as one of the complications in LCP repair.[12] [13] [14] The studies of Gallagher and colleagues[18] and Kääb and colleagues[22] demonstrate that off-axis insertion of locking screws with cross-threading significantly compromises mechanical stability to the locking plate–screw interface. Therefore, it is recommended to adequately insert and tighten locking screws to help maintain construct stiffness and to prevent locking screw unlocking.

The plate–bone construct setup in this study was designed to simulate a clinical comminuted fracture gap model in which the LCP system is commonly applied. A synthetic bone surrogate was used in this study to help standardize the application and testing of the implants by minimizing inter-specimen variation such as variance in size, breed, bone density, and conformation difference, which are common problems in cadaveric specimens. Although the bone surrogate used in this study has been validated to have mechanical properties close to native canine cortical bone,[31] [32] [33] previous studies indicate that the fourth-generation material of this bone model has superior screw-holding capacity compared with cadaveric bone.[31] [34] The increased screw-holding capacity could have reduced the possibility of screw failure within the bone surrogate in this study, leading to an underestimation of the failure mode. Since the aim of this study was the biomechanical comparison between different locking constructs, the authors believed the use of a synthetic bone model to be appropriate by limiting the confounding factors and premature failure of the cadaveric bone. However, the use of a synthetic bone model is one of the limitations of this study. Further research using cadaveric bone or less dense synthetic material would be beneficial as the mode of failure in this material may differ from that observed in the present study.

Due to the lack of a standardized specification for cyclic torsional testing, the testing methodology in the present study was designed based on previous biomechanical studies. The center of torsion was made along the axis of the bone surrogate as it best represents the rotational axis clinically. Centering torsion along the axis of the plate would likely favor the mechanical assessment of the plates rather than the construct as a whole unit. Both load and displacement control models are established testing methods for biomechanical assessment of implant constructs. Load control testing may more likely represent clinical scenarios. However, given that gait pattern is restricted during recovery and the current displacement control testing generated a consistent load that was less than the yield load of the constructs, the findings in this study still have good clinical relevance. The angular displacement range was chosen based on a previous study[12] and the estimated range of physiological angle in dogs.[10] [35] [36] The lack of testing under bending and axial compression in this study is a limitation.

Given the in vitro nature of the study model, there are other limitations as well. Due to the lack of a standardized testing protocol and similar studies comparing LCP constructs from different manufacturers, limited comparison of the present study results can be made. Additionally, the results of this study cannot be applied directly to clinical practice because in vivo factors were not created in this study model, but they are likely to vary clinically. There are also no consistent data or consensus on stiffness and cycles to failure required for LCP constructs until adequate bone healing occurs in a fracture repair. Nevertheless, we believe that the results of this study can still provide useful information for clinicians regarding the potential biomechanical differences between these implants. Further biomechanical studies under cyclic axial loading and bending as well as in vivo studies would be ideal for complete assessment.

Conclusion

This study reports the biomechanical difference in cyclic torsion of LCP constructs from DPS, KB, and Vi. Based on the results of the present study, the DPS constructs have superior torsional stiffness and cycles to failure compared with the Vi constructs, while there was no significant difference between the KB and Vi constructs. Although the constructs of DPS and KB showed comparable stiffness, the fatigue life of the DPS constructs was significantly longer. Further studies testing the construct under axial compression and bending, as well as clinical studies comparing fracture fixation outcome using these LCP implants, will provide additional data to guide clinical decision-making.

Conflict of Interest

D.R.J. is an occasional presenter and consultant on the products of Knight Benedikt. None of the other authors have any conflicts of interest to declare.

Acknowledgement

The authors would like to thank the Orthopaedic Biomechanics Research Group and the associated departments at Macquarie University for assistance with the mechanical setup and testing of this project. Special thanks to Dr. Vince Frank from Knight Benedikt for supporting this study. Knight Benedikt sponsored the implants and laboratory fees needed for the study.

Note

An abstract of this paper was presented at the Australian & New Zealand College of Veterinary Scientist (Surgery Chapter), Gold Coast, Australia, July 29, 2023.

Authors' Contribution

L.H.L. and D.R.J. contributed to conception of the study, study design, acquisition of data, and data analysis and interpretation. R.C.A. and J.C. contributed to the study design, acquisition of data, and data analysis and interpretation. All the authors drafted, revised, and approved the submitted manuscript.

• References

• 1 Hudson CC, Pozzi A, Lewis DD. Minimally invasive plate osteosynthesis: applications and techniques in dogs and cats. Vet Comp Orthop Traumatol 2009; 22 (03) 175-182
• 2 Gautier E, Sommer C. Guidelines for the clinical application of the LCP. Injury 2003; 34 (Suppl. 02) B63-B76
• 3 Stoffel K, Dieter U, Stachowiak G, Gächter A, Kuster MS. Biomechanical testing of the LCP–how can stability in locked internal fixators be controlled?. Injury 2003; 34 (Suppl. 02) B11-B19
• 4 Perren SM. Evolution of the internal fixation of long bone fractures. The scientific basis of biological internal fixation: choosing a new balance between stability and biology. J Bone Joint Surg Br 2002; 84 (08) 1093-1110
• 5 Carter DR, Vasu R, Spengler DM, Dueland RT. Stress fields in the unplated and plated canine femur calculated from in vivo strain measurements. J Biomech 1981; 14 (01) 63-70
• 6 Carter DR, Smith DJ, Spengler DM, Daly CH, Frankel VH. Measurement and analysis of in vivo bone strains on the canine radius and ulna. J Biomech 1980; 13 (01) 27-38
• 7 Coleman JC, Hart RT, Burr DB. Reconstructed bone end loads on the canine forelimb during gait. J Biomech 2003; 36 (12) 1837-1844
• 8 Coleman JC, Hart RT, Owan I, Tankano Y, Burr DB. Characterization of dynamic three-dimensional strain fields in the canine radius. J Biomech 2002; 35 (12) 1677-1683
• 9 Gautier E, Perren SM, Cordey J. Strain distribution in plated and unplated sheep tibia an in vivo experiment. Injury 2000; 31 (Suppl. 03) C37-C44
• 10 Page AE, Allan C, Jasty M, Harrigan TP, Bragdon CR, Harris WH. Determination of loading parameters in the canine hip in vivo. J Biomech 1993; 26 (4–5): 571-579
• 11 Shahar R, Banks-Sills L, Eliasy R. Stress and strain distribution in the intact canine femur: finite element analysis. Med Eng Phys 2003; 25 (05) 387-395
• 12 Bilmont A, Palierne S, Verset M, Swider P, Autefage A. Biomechanical comparison of two locking plate constructs under cyclic torsional loading in a fracture gap model. Two screws versus three screws per fragment. Vet Comp Orthop Traumatol 2015; 28 (05) 323-330
• 13 Filipowicz D, Lanz O, McLaughlin R, Elder S, Werre S. A biomechanical comparison of 3.5 locking compression plate fixation to 3.5 limited contact dynamic compression plate fixation in a canine cadaveric distal humeral metaphyseal gap model. Vet Comp Orthop Traumatol 2009; 22 (04) 270-277
• 14 Ahmad M, Nanda R, Bajwa AS, Candal-Couto J, Green S, Hui AC. Biomechanical testing of the locking compression plate: when does the distance between bone and implant significantly reduce construct stability?. Injury 2007; 38 (03) 358-364
• 15 Fitzpatrick DC, Doornink J, Madey SM, Bottlang M. Relative stability of conventional and locked plating fixation in a model of the osteoporotic femoral diaphysis. Clin Biomech (Bristol, Avon) 2009; 24 (02) 203-209
• 16 Kanchanomai C, Phiphobmongkol V, Muanjan P. Fatigue failure of an orthopedic implant: a locking compression plate. Eng Fail Anal 2008; 15 (05) 521-530
• 17 Kaczmarek J, Bartkowiak T, Paczos P, Gapinski B, Jader H, Unger M. How do the locking screws lock? A micro-CT study of 3.5-mm locking screw mechanism. Vet Comp Orthop Traumatol 2020; 33 (05) 316-326
• 18 Gallagher B, Silva MJ, Ricci WM. Effect of off-axis screw insertion, insertion torque, and plate contouring on locked screw strength. J Orthop Trauma 2014; 28 (07) 427-432
• 19 Boudreau B, Benamou J, von Pfeil DJF, Guillou RP, Beckett C, Déjardin LM. Effect of screw insertion torque on mechanical properties of four locking systems. Vet Surg 2013; 42 (05) 535-543
• 20 Pearson T, Glyde MR, Day RE, Hosgood GL. The effect of intramedullary pin size and plate working length on plate strain in locking compression plate-rod constructs under axial load. Vet Comp Orthop Traumatol 2016; 29 (06) 451-458
• 21 Stoffel K, Lorenz KU, Kuster MS. Biomechanical considerations in plate osteosynthesis: the effect of plate-to-bone compression with and without angular screw stability. J Orthop Trauma 2007; 21 (06) 362-368
• 22 Kääb MJ, Frenk A, Schmeling A, Schaser K, Schütz M, Haas NP. Locked internal fixator: sensitivity of screw/plate stability to the correct insertion angle of the screw. J Orthop Trauma 2004; 18 (08) 483-487
• 23 Palierne S, Blondel M, Swider P, Autefage A. Biomechanical comparison of use of two screws versus three screws per fragment with locking plate constructs under cyclic loading in compression in a fracture gap model. Vet Comp Orthop Traumatol 2022; 35 (03) 166-174
• 24 Palierne S, Froidefond B, Swider P, Autefage A. Biomechanical comparison of two locking plate constructs under cyclic loading in four-point bending in a fracture gap model: two screws versus three screws per fragment. Vet Comp Orthop Traumatol 2019; 32 (01) 59-66
• 25 Chao P, Conrad BP, Lewis DD, Horodyski M, Pozzi A. Effect of plate working length on plate stiffness and cyclic fatigue life in a cadaveric femoral fracture gap model stabilized with a 12-hole 2.4 mm locking compression plate. BMC Vet Res 2013; 9: 125
• 26 Blake CA, Boudrieau RJ, Torrance BS. et al. Single cycle to failure in bending of three standard and five locking plates and plate constructs. Vet Comp Orthop Traumatol 2011; 24 (06) 408-417
• 27 Cabassu JB, Kowaleski MP, Skorinko JK, Blake CA, Gaudette GR, Boudrieau RJ. Single cycle to failure in torsion of three standard and five locking plate constructs. Vet Comp Orthop Traumatol 2011; 24 (06) 418-425
• 28 DeTora M, Kraus K. Mechanical testing of 3.5 mm locking and non-locking bone plates. Vet Comp Orthop Traumatol 2008; 21 (04) 318-322
• 29 Aguila AZ, Manos JM, Orlansky AS, Todhunter RJ, Trotter EJ, Van der Meulen MCH. In vitro biomechanical comparison of limited contat dynamic compression plate and locking compression plate. Vet Comp Orthop Traumatol 2005; 18 (04) 220-226
• 30 Roe S. Biomechanics of fracture fixation. Vet Clin North Am Small Anim Pract 2020; 50 (01) 1-15
• 31 Zdero R, Brzozowski P, Schemitsch EH. Biomechanical properties of artificial bones made by Sawbones: a review. Med Eng Phys 2023; 118: 104017
• 32 Heiner AD. Structural properties of fourth-generation composite femurs and tibias. J Biomech 2008; 41 (15) 3282-3284
• 33 Chong ACM, Miller F, Buxton M, Friis EA. Fracture toughness and fatigue crack propagation rate of short fiber reinforced epoxy composites for analogue cortical bone. J Biomech Eng 2007; 129 (04) 487-493
• 34 Aziz MSR, Nicayenzi B, Crookshank MC, Bougherara H, Schemitsch EH, Zdero R. Biomechanical measurements of cortical screw purchase in five types of human and artificial humeri. J Mech Behav Biomed Mater 2014; 30: 159-167
• 35 Fischer MS, Lehmann SV, Andrada E. Three-dimensional kinematics of canine hind limbs: in vivo, biplanar, high-frequency fluoroscopic analysis of four breeds during walking and trotting. Sci Rep 2018; 8 (01) 16982
• 36 Ellis RG, Rankin JW, Hutchinson JR. Limb kinematics, kinetics and muscle dynamics during the sit-to-stand transition in greyhounds. Front Bioeng Biotechnol 2018; 6: 162

Publication History

Received: 08 October 2023

Accepted: 18 July 2024

Article published online:
05 August 2024