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. 2013 Nov 12;472(4):1310–1317. doi: 10.1007/s11999-013-3379-7

Dual Plating of Humeral Shaft Fractures: Orthogonal Plates Biomechanically Outperform Side-by-Side Plates

Victor Kosmopoulos 1,✉, Arvind D Nana 1,2
PMCID: PMC3940765  PMID: 24218163

Abstract

Background

Single large-fragment plate constructs currently are the norm for internal fixation of middiaphyseal humerus fractures. In cases where humeral size is limited, however, dual small-fragment locking plate constructs may serve as an alternative. The mechanical effects of different possible plate configurations around the humeral diaphysis may be important, but to our knowledge, have yet to be investigated.

Questions/purposes

We used finite element analysis to compare the simulated mechanical performance of five different dual small-fragment locking plate construct configurations for humeral middiaphyseal fracture fixation in terms of (1) stiffness, (2) stress shielding of bone, (3) hardware stresses, and (4) interfragmentary strain.

Methods

Middiaphyseal humeral fracture fixation was simulated using the finite element method. Three 90° and two side-by-side seven-hole and nine-hole small-fragment dual locking plate configurations were tested in compression, torsion, and combined loading. The configurations chosen are based on implantation using either a posterior or anterolateral approach.

Results

All three of the 90° configurations were more effective in restoring the intact compressive and torsional stiffness as compared with the side-by-side configurations, resulted in less stress shielding and stressed hardware, and showed interfragmentary strains between 5% to 10% in torsion and combined loading.

Conclusions

The nine-hole plate anterior and seven-hole plate lateral (90° apart) configuration provided the best fixation. Our findings show the mechanical importance of plate placement with relation to loading in dual-plate fracture-fixation constructs.

Clinical Relevance

The results presented provide novel biomechanical information for the orthopaedic surgeon considering different treatment options for middiaphyseal humeral fractures.

Introduction

Locking plates for fracture fixation offer biologic and mechanical advantages and therefore have increased in popularity. Locking the screws to the plate allows the plate to sit at a distance offset from the underlying bone surface providing a biologic advantage for bone fracture healing by preserving the periosteal blood supply underlying the plate [12, 13, 26, 31]. Mechanically this provides stability without the need for the plate to match the curvature of the bone surface and without the need to compress and maintain friction between the plate and bone surface [4, 7, 31].

Currently, the majority of diaphyseal humeral fractures are treated nonoperatively. In the subset treated operatively, one large-fragment plate construct is considered the norm for internal fixation. The smaller size of the humerus in some patients, however, limits the diaphyseal shaft length and/or diameter available for fixation, and therefore makes using a large-fragment plate difficult. Difficulties that arise include the number of screws that can be placed, a resulting bulky fixation with undesirable stress shielding [19], and having to precontour the large-fragment plate to match the diverging anatomy of the humeral metaphysis.

Recent literature suggests that dual small-fragment plating constructs may be mechanically superior to one large-fragment plate construct [32] and may have a role in the fixation of certain fracture patterns. Furthermore, small-fragment plates for humeral shaft fracture fixation have shown promising clinical results [18]. Questions remain, however, regarding how plate placement on the diaphysis affects the mechanical performance of the dual-plate construct, necessitating further research. Plates typically are implanted using either a posterior or anterolateral approach. The posterior approach allows for direct observation of the fracture and posterior and lateral plate placement but requires the nerve to be dissected out because it is in the middle of the operative field. In contrast, the anterolateral approach avoids direct observation of the nerve and allows for anterior and lateral plate placement.

We used a finite element (FE) method, a computational simulation approach, to compare the mechanical performance of five different dual small-fragment locking plate constructs for humeral shaft fracture fixation under three different loading regimens. The five dual-plate placement configurations chosen are based on implantation using the aforementioned surgical approaches. Using the different constructs and the three loading regimens (eccentric compression, torsion, and combined) in a FE humerus fracture model, we aimed to determine the best plate configuration by assessing (1) stiffness, (2) bone stress shielding, (3) hardware stress, and (4) interfragmentary strain.

Materials and Methods

Six (one intact and five fixation) FE models were created and analyzed using COMSOL Multiphysics® (Version 4.3; COMSOL, Inc, Burlington, MA, USA). The intact humeral cortical bone diaphysis was idealized as a hollow cylinder with a constant cross section. Dimensions were based on reported mean values from an MRI study of 20 volunteers with an average age of 37 years [16]. The length of the tested portion of the diaphysis was set to 237 mm and the diameters of the canal and outer bone cortex were 12.1 mm and 19.3 mm, respectively [16]. Cortical bone was modeled as an orthotropic material characterized by nine independent technical constants (E1 = 12.0, E2 = 13.4, E3 = 20.0 GPa; G12 = 4.53, G23 = 6.23, G13 = 5.61 GPa; ν12 = 0.376; ν23 = 0.235, ν13 = 0.222) [2, 3] with density set to 1817 kg/m3 [22].

To study a worst-case scenario such as an unstable comminuted fracture, a 1-cm transverse fracture gap was created at the middiaphysis of the intact model (simulating an Orthopaedic Trauma Association 12.C.2 fracture) [1, 5, 6, 18, 19, 23, 30, 32]. Fracture fixation was performed using five different seven-hole and nine-hole small-fragment (3.5 mm) locking-plate configurations as follows: (1) nine-hole plate anterior and seven-hole plate lateral, 90° apart (Fig. 1A); (2) nine-hole plate lateral and seven-hole plate anterior, 90° apart (Fig. 1B); (3) nine-hole plate anterolateral and seven-hole plate posterolateral, 90° apart (Fig. 1C); (4) nine-hole plate anterolateral and seven-hole plate posterolateral, side by side 65° apart (Fig. 1D); and (5) nine-hole plate posterolateral and seven-hole plate posteromedial, side by side 65° apart (Fig. 1E). The plates in all of the aforementioned models were centered to the fracture gap and locked at a 1-mm offset distance from the outer cortical bone cortex [1, 5, 25]. The plates are fixed using a total of eight bicortical locking screws [18, 25, 31] sequentially alternating between the two plates (four screws/plate) resulting in all the models having the same number of fixation points and equal working lengths. Such a staggered placement of the hardware has been suggested to reduce the risk of plate fracture [17]. Bicortical screws were chosen to stabilize the screw-bone interface [6, 7]. The generic locking plates and screws were modeled using stainless steel material properties (E = 193.0 GPa; ν = 0.3; ρ = 8000 kg/m3) [24]. Geometric properties of the seven-hole and nine-hole small-fragment locking plates and screws are similar to FDA-approved commercial hardware (Table 1).

Fig. 1A–E.

Fig. 1A–E

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The (A) nine-hole plate anterior and seven-hole plate lateral, 90° apart (Model A); (B) nine-hole plate lateral and seven-hole plate anterior, 90° apart (Model B); (C) nine-hole plate anterolateral and seven-hole plate posterolateral, 90° apart (Model C); (D) nine-hole plate anterolateral and seven-hole plate posterolateral, side by side 65° apart (Model D); and (E) nine-hole plate posterolateral and seven-hole plate posteromedial, side by side 65° apart (Model E) fixation construct configurations are shown. The humeral diaphysis is light gray and the plates and screws are dark gray. Lateral and anterior directions are oriented to the left and bottom, respectively.

Table 1.

Geometric properties of the small-fragment plates and screws used in Models A to E

Description Length (mm) Width (mm) Hole spacing (mm) Number of holes
3.5-mm small-fragment 9-hole plate 140 11 14.5 9
3.5-mm small-fragment 7-hole plate 111 11 14.5 7
Core length (mm) Core diameter (mm) Head height (mm) Head diameter (mm)
Locking screw 22 2.7 3.2 6.8
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Three different loading conditions were applied to test the fixation constructs [8, 11]: (1) eccentric compression; (2) torsion; and (3) combined eccentric compression and torsion (Fig. 2). The eccentric load was inferiorly directed and applied in 20-N increments to a maximum of 100 N a distance 40 mm posteromedial to the central longitudinal humeral axis at a rotation of 23.3° from the frontal plane [10]. This off-center eccentric loading produces combined bending and compressive loads on the humeral diaphysis. Torsion was applied in 1.0-Nm increments from 0.5 Nm to a maximum of 4.5 Nm along the central longitudinal axis of the humerus. Combined loading included simultaneous application of the eccentric compressive and torsional loads (eg, first load step = 20 N compression and 0.5 Nm torsion; last load step = 100 N compression and 4.5 Nm torsion).

Fig. 2A–C.

Fig. 2A–C

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The nine-hole plate anterior and seven-hole plate lateral, 90° configuration (Model A) is used to illustrate the simulated (A) eccentric compression, (B) torsion, and (C) combined loading conditions. The force (F) was applied in 20-N increments to a maximum of 100 N. Torsion was applied in 1-Nm increments from 0.5 Nm to a maximum of 4.5 Nm. The superior (Sup) and inferior (Inf) cross-sectional bone surfaces used to calculate the interfragmentary strain also are shown.

All outcomes, other than stiffness, are reported at the maximum loads simulated. The extrinsic compressive stiffness for each simulation was calculated as the slope of the applied compressive load to deformation curve. An analogous rotational stiffness was calculated as the slope of the applied torsional load to the resulting rotational deformation. Stiffness of each of the fixation constructs was compared with the intact humerus stiffness for each respective loading regimen with positive percentile differences denoting increases and negative percentile differences denoting decreases in stiffness. Von Mises stress distributions were computed for the bone, screws, and plates. Average bone stress analysis results for each of the fixation constructs were compared with the intact humerus model. In this manner, negative percentile changes denote decreased bone tissue stresses as compared with the intact humerus (indicative of bone stress shielding after fixation). To characterize the load-sharing performance of the construct for the screws, two additional measures were calculated. The first measure, denoted as the maximum-minimum range, is calculated as the difference in mean stress between the highest and lowest stressed screws. The second, the screw-to-screw fluctuation, is calculated as the difference in mean stress between adjacent screws. Lower values in these measures represent better load-sharing characteristics among the screws. Finally, the interfragmentary strain was characterized by dividing the interfragmentary motion by the fracture gap size. The interfragmentary motion for each construct was calculated by adding the maximum three-dimensional displacement of the cross-sectional bone surface superior and cross-sectional bone surface inferior to the fracture (Fig. 2).

Results

All three of the 90° configurations (Models A through C) were nearly equally effective in restoring the intact compressive stiffness showing less than a 2% difference in compression and combined loading, respectively (Fig. 3A). As a result of geometric equivalency in torsion, Models A through C behaved equally as did Models D and E. Models A through C were more effective however, in restoring the intact torsional stiffness (Fig. 3B). Model E, with the plates placed side by side posteriorly, was the only construct to exceed the intact compressive stiffness.

Fig. 3A–B.

Fig. 3A–B

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The relative stiffness of each of the tested constructs as a percentage of (A) the intact compressive (317 N/mm) and combined compressive (293 N/mm) stiffness, and (B) the intact torsional (5 Nm/degree) and combined torsional (5 Nm/degree) stiffness are shown. Positive percentile differences denote increases and negative percentile differences denote decreases as compared with the intact stiffness.

All of the fixation models resulted in some degree of stress shielding by redistributing the load and consequently reducing the average stress on the humerus (Fig. 4). The side-by-side configuration from Model D showed the least stress shielding in compression, whereas Models A through C, with the 90° configuration, were better in torsion and combined loading.

Fig. 4.

Fig. 4

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The percentages of change in mean von Mises bone stress as compared with the intact humeral stress in compression (3.4 MPa), torsion (5.4 MPa), and combined loading (6.7 MPa) are shown. The results are shown for each of the five construct configurations tested (Models A through E) in maximum compression (100 N), torsion (4.5 Nm), and combined (100 N, 4.5 Nm) loading. In torsion, as a result of geometric equivalency, results are identical among Models A to C and between Models D and E.

The fixation plates in Models A through C were, on average, less stressed in torsion and combined loading than the plates from the side-by-side constructs (Models D and E). Model E however, with the posteriorly placed plates, was significantly more effective in reducing plate stresses in compression (Fig. 5). The highest stress concentrations were located near and around the unused screw holes for each plate and at the neck of the screws just below the plates for all the constructs studied (Fig. 6). Generally, as evidenced by the lower mean screw stresses, Models A through C were better at reducing screw stress in torsion and combined loading than the side-by-side plate constructs (Models D and E) (Table 2). Using the screw-to-screw fluctuation and maximum-minimum range measures however, Model E showed the best load-sharing characteristics between screws in combined loading.

Fig. 5.

Fig. 5

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The mean von Mises plate stress comparisons among each of the five construct configurations tested (Models A through E) in maximum compression (100 N), torsion (4.5 Nm), and combined (100 N, 4.5 Nm) loading are shown. In torsion, as a result of geometric equivalency, the results are identical among Models A through C and between Models D and E.

Fig. 6A–B.

Fig. 6A–B

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The von Mises hardware stress distributions at the maximum combined load simulated (4.5 Nm torsion, 100 N compression) for the best-performing configuration with the (A) nine-hole plate anterior and seven-hole plate lateral, 90° apart (Model A); and worst-performing configuration with the (B) nine-hole plate anterolateral and seven-hole plate posterolateral, side by side 65° apart (Model D). Increasing element stress magnitudes are illustrated from red to blue in the color bar legend and scaled to a maximum of 300 MPa to allow for direct visual comparison.

Table 2.

Mean screw stress, fluctuation, maximum-minimum range, and locations for the highest stressed screws*

Loading Model
A B C D E
Compression
 Mean (MPa) 9.5 9.6 9.6 11.6 6.7
 Screw-to-screw (MPa) 1.5 1.6 1.6 2.5 2.5
 Maximum-minimum (MPa) 5.4 5.5 5.4 8.8 4.4
 Screw locations with highest stress 3, 6 4, 5 4, 5 4, 5 2, 7
Torsion†
 Mean (MPa) 37.4 37.4 37.4 42.7 42.7
 Screw-to-screw (MPa) 6.3 6.3 6.3 7.9 7.9
 Maximum-minimum (MPa) 22.1 22.1 22.1 27.9 27.9
 Screw locations with highest stress 4, 5 4, 5 4, 5 4, 5 4, 5
Combined
 Mean (MPa) 38.9 38.9 38.9 44.7 43.8
 Screw-to-screw (MPa) 8.9 8.4 8.3 10.5 7.8
 Maximum-minimum (MPa) 30.4 30.2 30.2 37.5 27.7
 Screw locations with highest stress 4, 6 3, 5 4, 6 4, 6 4, 5
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* Details in Materials and Methods; †because of geometric equivalency in torsion, the measures for Models A through C and for Models D and E were equal.

The 90° configurations tested in Models B and C resulted in the highest (9.4%) and lowest (7.4%) interfragmentary strains in combined loading, respectively (Fig. 7). In torsion, the 90° constructs (Models A through C) resulted in lower interfragmentary strains than the side-by-side constructs (Models D and E). Models D and E however, showed the highest (5.0%) and lowest (4.4%) strain at the fracture site in compression, respectively.

Fig. 7.

Fig. 7

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The results for percent interfragmentary strain for the five construct configurations tested (Models A through E) are shown. The results are reported at the maximum compressive (100 N), torsional (4.5 Nm), and combined (100 N, 4.5 Nm) loads. In torsion, as a result of geometric equivalency, the results are identical among Models A through C and between Models D and E.

Discussion

The smaller size of the humerus in some patients may limit application of the more commonly used large-fragment plate constructs. In such cases, large-fragment plates may limit the number of screws that can be inserted (ie, less holes/unit length), and lead to increased stress shielding from the greater mismatch in load transfer between bone and plate [19, 32]. Dual small-fragment plate constructs may offer a promising alternative. Dual 3.5-mm locking plates offer advantages over one large fragment plate including: (1) the 3.5 mm plate is more easily contoured; (2) the 3.5 mm plate width easily accommodates small bone diameters; and (3) dual 3.5 mm plates require a smaller incision and working length compared with one large fragment plate. The biomechanical benefits related to different placement configurations of the two plates, however, remain unanswered. Using FE modeling, we compared the performance of five different small-fragment dual-plate configurations for fixation of middiaphyseal humeral fractures by evaluating (1) stiffness, (2) bone stress shielding, (3) hardware stresses, and (4) interfragmentary strain.

Limitations

The results presented are based on the specific plates and screws modeled and may not be representative for other plates and screws. Nevertheless, the tested plates and screws were sized to be similar to what currently is commercially available. Additionally, the use of FE modeling offers the advantage of studying bone stress shielding as an outcome measure. Similar to an experimental cadaveric study, the computational simulations presented do not evaluate the in vivo bone remodeling response expected after internal fixation. Clinically, the mechanics of fixation are expected to change with stress shielding (eg, bone resorption) or bone fracture healing, especially with union of the fracture gap. Fatigue, like with remodeling, is another time-related response that was not simulated in this study. Although fatigue microdamage and crack propagation would provide additional valuable information, it also would add to the complexity of this numerical study and outcome variables considered [14]. Screw fixation, for example, was fixed in the models presented. Clinically, after cyclic loading, the behavior at the screw-bone interface may weaken [9]. Even with these limitations, however, the comparisons as presented provide worthwhile and novel information for the orthopaedic surgeon considering different treatment options. The stress results from the study, for example, detail the changes in load transfer for the bone and each part of the construct otherwise difficult to measure clinically and experimentally.

Construct stiffness is of high importance. It governs the performance of the fixation system as indicated, for example using the other interrelated outcomes measures reported (bone stress shielding, hardware stress, and interfragmentary strain). Excessive stiffness of the fixation construct reduces stress and strain on the bone and may lead to bone resorption [13, 28] and in time screw loosening and construct failure [4, 7, 15, 25]. However, excessive reductions in stiffness may lead to increasing screw and plate stress and early fatigue failure of the construct [7, 25, 26]. Stiffness resulting from the compressive loading simulations of the humerus may be especially important during crutch weightbearing [18, 19]. Moreover, torsional loading also is of interest in the analysis of humeral fracture fixation constructs because it has been reported as a predominant loading mode and possible cause for nonunion of humeral fractures [7, 8, 29, 30]. In compression and torsion, and consequently combined loading, the 90° configurations (Models A through C) were found to outperform the side-by-side constructs in stiffness recovery as compared with the intact humerus. In our simulations, the stiffness was relatively insensitive to exchanging the seven-hole and nine-hole plates, as is done between Models A and B, or rotating the 90° configuration, as is done between Models A and C. In contrast, the side-by-side posterior placement of the plates (Model E), closer to the site of compressive load application, reduced bending loads on the plates and increased the construct stiffness beyond that of the intact humerus.

Ideally, when loaded, the fixation construct design should be balanced by reducing bone stress shielding yet maintaining adequate fixation. Model E, with the highest compressive stiffness, also had the highest levels of average stress shielding for all three of the loading conditions tested. Relative to each other however, all the model configurations tested resulted in less than a 5% difference in stress shielding. Although in compression and combined loading the tested configurations resulted in an approximately 40% to 45% reduction in bone tissue stress, this was not the case in torsion. The configurations were less effective in shielding bone shear stress. As recommended in the literature [1, 5, 25], all of the locking plates modeled were offset 1 mm from the cortex avoiding undue stress shielding and contact below the plate. This advantage with locked plates has been suggested to prevent local bone necrosis [20].

The stability of the fixation system is influenced by hardware factors including the number of screws, type of screws (ie, bicortical, unicortical), working length, plate offset from the bone cortex, and placement of the hardware [1, 5–7, 18, 25, 31]. Other than plate placement, these aforementioned hardware variables were controlled in this study based on recommendations from the literature. Four screws per fragment were used based on findings that additional screws did not show a significant increase in torsional stiffness [25]. In gauging performance, a goal of the fixation system should be to reduce and more evenly distribute the applied stress among the hardware components of the construct [27]. This in turn will help limit stress risers and extend the fatigue life and strength of the system. O’Toole et al. [18], comparing single-plate locking and nonlocking 3.5-mm small-fragment constructs for humeral shaft fixation, reported both constructs withstood strenuous fatigue and axially failed above anticipated physiologic loads. Results from all of the dual-plate locking constructs compared in our study show the highest stress concentrations occurring at the neck of the screws just below the screw head and plate. These modeling results confirm clinical findings suggesting this may be the most likely location for hardware failure [27, 29]. This is especially true for the highest stressed screws which varied in location based on plate configuration and type of loading. In compression, Models A and E generally outperformed the other constructs by more evenly distributing the applied load among the two plates, with the larger nine-hole plates sustaining a slightly greater portion of the stress. As evidenced by the lower mean screw stress, screw-to-screw fluctuations, and maximum-minimum ranges, these models also outperformed the other constructs with respect to load sharing between screws. However, the highest stressed screws in Model A were located adjacent to the fracture on the nine-hole plate. In contrast, the highest stressed screws in Model E were at the far ends of the seven-hole plate. This is indicative of the greater bending loads on Model A and the greater axial loads on Model E. Thus, understanding how the plate will be loaded based on its placement is important to reduce hardware stress and improve the load-sharing characteristics of the construct. This may help prolong and/or prevent screw failure.

Fixation with locked plate constructs aims to minimize motion while tolerating an increased fracture gap [4]. Some motion at the fracture site is favored to promote secondary bone healing [4, 21, 27]. Secondary bone healing has been reported to occur when interfragmentary strain is kept between 2% to 10% [4]. Although biologic responses were not addressed in our study, all of the tested configurations did satisfy this interfragmentary strain criteria for all the loading regimens simulated. Strain at the fracture site was highest (exceeding 7%) for all the models when combined loading was applied.

Based on the simulations performed and relative comparison between outcomes from this study, the 90° configuration with a nine-hole plate placed anteriorly and a seven-hole plate placed laterally (Model A) was found to mechanically outperform the side-by-side constructs and slightly outperform the other 90° configurations studied. As loading of the humerus in vivo is likely a combination of compression and torsion, this configuration was one of the most effective in restoring the intact stiffness and reducing bone stress shielding and hardware stress while meeting the suggested interfragmentary strain criteria. Future studies are needed comparing large-fragment plates, nonlocking small-fragment plates, and nail fixation to the proposed dual small-fragment constructs to help establish optimal fixation for middiaphyseal humeral fractures. Although further clinical studies are needed to confirm our main findings, the mechanical findings presented using dual 3.5-mm small-fragment locking plate configurations are promising.

Acknowledgments

We thank the University of North Texas Health Science Center for providing the computational resources needed to complete this study.

Footnotes

Each author certifies that he or she, or a member of his or her immediate family, has no funding or commercial associations (eg, consultancies, stock ownership, equity interest, patent/licensing arrangements, etc) that might pose a conflict of interest in connection with the submitted article.

All ICMJE Conflict of Interest Forms for authors and Clinical Orthopaedics and Related Research editors and board members are on file with the publication and can be viewed on request.

Clinical Orthopaedics and Related Research neither advocates nor endorses the use of any treatment, drug, or device. Readers are encouraged to always seek additional information, including FDA-approval status, of any drug or device prior to clinical use.

This work was performed at the University of North Texas Health Science Center, Fort Worth, TX, USA.

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