Three-Dimensional Printing of Surgical Clips: An In Vitro Pilot Study and Trial of Efficacy

Introduction

The past few years have brought much attention to the potential of three-dimensional (3D) printing, with significant coverage in the national press. ¹ Also known as additive manufacturing because of the way materials are printed layer by layer, 3D printing was invented in the 1980s, but has seen expansive growth over the past decade. With the ability to render almost any computer-aided design (CAD) drawing into a 3D object, 3D printing has potentially more diverse roles than traditional manufacturing modalities.

This technology has already been utilized in numerous fields, including medicine, with potential benefits of cost savings with manufacturing, decreased packaging and transport to limit environmental impact, broader availability of technology, and ability to personalize equipment. In surgery, specifically, the development of 3D printed surgical instruments ^{2 –4} and permanent implants has already been reported. ⁵ And in urology, 3D printing has been utilized to build laparoscopic trocars, ⁶ surgical planning models, ⁷ and ureteral stents. ⁶

One of the more common utilized devices in minimally invasive urologic surgery is the Weck^® Hem-o-lok^® clip (Teleflex, Inc., Wayne, PA). A nonabsorbable polymer, the clip has become commonplace for large blood vessel ligation during nephrectomy, ⁸ athermal dissection during radical prostatectomy, and renorrhaphy during partial nephrectomy. ⁹ Our goal was to design a surgical clip using the Hem-o-lok as a model, and use 3D printing technology to produce a working prototype.

Materials and Methods

Our experimental clip CAD was reverse engineered with SolidWorks (Waltham, MA) using a 10 mm Hem-o-lok polymer clip as a model. Printing was done using an Objet Connex500 multijetting system (Stratasys, Eden Prairie, MN), which allowed the use and adjustment of multiple proprietary polymers to optimize design properties.

Experimental design

Two aspects were tested with regard to our final printed clip design. First, we quantified the ability of the printed clips to lock appropriately without fracturing using the clip applier, by closing the clips across a ¼″ width Penrose drain. Trials were also performed with commercial Hem-o-lok clips as controls. Second, we tested the closing pressure of the printed clips by filling the Penrose drain with normal saline. The saline was infused through a pressure bag to slowly increase filling pressure until the clip leaked or the Penrose ruptured. Leak pressures were measured using a digital pressure gauge (Fig. 1). The experiment was also repeated with model clips as controls.

OPEN IN VIEWER

FIG. 1.

3D printed clip (top) compared with manufactured clip (bottom) placed in Hem-o-lok^® clip applier. 3D = three-dimensional.

Results

Initial clips were printed in multiples of 50 using only Objet VeroWhitePlus (RGD835), a polymer that became rigid and held shape well after curing. Testing demonstrated that clips would invariably fracture at the hinge upon closure to itself, without the addition of a Penrose drain. Ultimately, the addition of Objet TangoBlackPlus (FLX980) (a rubber-like polymer) at the hinge was utilized to improve flexibility. Multiple copies of the same clip were printed with varying hinge mixes until the optimal combination of flexibility and stiffness was identified at a shore value of 70. Final adjustments in the clip design were made to allow functionality with the Weck Hem-o-lok clip applier (Fig. 2).

OPEN IN VIEWER

FIG. 2.

Experimental setup showing fluid tubing, digital pressure gauge, and clip placed on end of penrose.

After adjustment of the material to TangoBlackPlus in the vicinity of the live hinge, 3D printed clips closed appropriately upon themselves without the addition of a Penrose drain. Subsequent to this, the Penrose drain was introduced in our ex vivo experimental setup. In total, 50 3D printed surgical clips and 30 commercial clips were tested for locking ability over the Penrose. Fracture rate during closure of the printed clips was 54% (n = 27), whereas none of the Hem-o-lok clips broke (0%, p < 0.0001). Printed clips were noted to always fracture at the locking tabs, whereas the hinge did demonstrate appropriate flexibility without fracturing.

Of the 23 printed clips that closed effectively, pressure testing demonstrated fluid leakage at a mean 20.7 ± 11.7 κPa (155.1 mm Hg). None of the 30 commercial clips leaked fluid until the burst of the Penrose at mean 46.2 ± 0.7 κPa (346.5 mm Hg, p < 0.0001). Using 220 mm Hg (29.3 κPa) as an arbitrary cutoff for surgical use, six printed surgical clips (12%) would have achieved success.

Discussion

We demonstrate the feasibility of design, production, and testing of 3D printed surgical clips based on a Hem-o-lok clip model. Although the initial design was not functional because of rigid hinge material, the ability to selectively adjust polymers in different areas of the printed clip resulted in improved hinge flexibility. However, even with these capabilities, our 3D printed clips still demonstrated significant fragility as more than half fractured at the clasp during closure without placement on tissue. Although a rigid material like VeroWhitePlus was needed to support the clasp, the properties of this specific polymer still demonstrate potential limitations with this technology, depending on the printer utilized and the materials available. In addition, only 12% of clips (26% of those that did not fracture) actually demonstrated appropriate theoretical closing pressure (220 mm Hg) to be utilized during surgery. In contrast, standard Hem-o-lok clips all locked close without tissue in between, and all held significant closing pressure throughout testing, not an unexpected result given the superior material characteristics of plastic injection molding.

There has been growing interest in 3D printing of surgical tools and materials for intraoperative use. For example, the idea that on-demand printing could minimize storage and packaging requirements has prompted advancement in the military. In 2009, the National Institute of Health released a Small Business Innovation Research/Small Business Technology Transfer grant request that included additive manufacturing of on-demand surgical instruments (PA-09-113). The Defense Advanced Research Projects Agency pursued the goal of designing and printing an on-demand surgical kit for battlefield trauma, given the potential limitations of storage and transport in times of conflict. ³ Using fused-deposition modeling (FDM), they developed a hemostat, needle driver, forceps, metzenbaum scissors, retractor, and scalpel, made of acrylonitrile butadiene styrene (ABS) plastic. They noted a 6-hour print time for the entire set, and all instruments functioned well during trials on a chicken carcass and a simulator. With a similar pursuit, Rankin et al. published their work designing and printing functional Army/Navy retractors. ² They were able to effectively print retractors in <90 minutes, and their polylactic acid product was sterile. Their retractors passed anticipated strength testing, with an added benefit of costing only an estimated $0.46 of material.

Similar advancements have been pursued in the aerospace industry. In 2014, a printer produced by Made in Space (Mountain View, CA) effectively printed on board the international space station. That same year, Wong and Pfahnl published their work by designing a 3D printed surgical kit using ABS plastic. ⁴ With the goal of designing a surgical kit to be used during space missions, they were able to print 10 different instruments including forceps, hemostats, clamps, scalpel handles, and sponge sticks. Although printing times ranged from 50 minutes to 10.5 hours per instrument, multiple surgeon testing proved that the instruments would likely perform adequately. Although the pursuit of military and space development is different than daily surgical needs, the same benefits of limiting packaging, transport, and storage can apply.

In urology, specifically, development of 3D printed surgical tools has been of interest. Del Junco and colleagues effectively printed both functional ureteral stents and laparoscopic trocars. ⁶ Their ureteral stents were initially printed with an Objet Connex500 material jetting printer, similar to our study. However, production of the internal lumen required printing of a dissolvable support material, which limited product size. Therefore, the smallest functional stent they initially printed was 12F. They subsequently utilized a powder bed fusion printer (Eosint P395; EOS e-Manufacturing Solutions, Krailing, Germany), which laser prints into a powder base that provides its own internal support during printing. They were able to print a functional nylon 9F ureteral stent, with an internal diameter that accepted a standard 0.035″ guidewire, and effectively passed the stent into both a cadaver and porcine ureter. The same group also demonstrated that their 3D printed stents had flow rates comparable with those of commercial stents in an ex vivo porcine model. ¹⁰ Although the design, production, and testing of this stent were effective, most standard ureteral stents are in the 6F to 7F range. Therefore, similar to our work, printing and material refinement are likely required before utilization of such equipment becomes standard.

The same group attempted to employ this technology with laparoscopic trocars. They printed their 3D design with an Objet30 material printer, which took ∼5 hours for four trocars. ⁶ All printed trocars were larger than the model trocars they were based upon (outer diameter 9.40 mm vs 7.00 mm and 7.87 mm), with significantly larger skin defects (p < 0.001). Although the printed trocars accepted standard 5 mm laparoscopic instruments and maintained pneumoperitoneum of 15 mm Hg, this again highlights size limitations compared with standard device manufacturing.

Other work in 3D printing of urologic products includes an antireflux valve positioning on the distal end of a ureteral stent. ¹¹ Park and colleagues printed the valve with an Objet300 material jetting printer. In comparison with conventional double pigtail stents, they demonstrated that their 3D printed flap valve maintained similar antegrade flow, but significantly reduced retrograde flow.

There are multiple limitations that deserve mention regarding 3D printing of surgical products. First, the safety and regulation of printed products could be difficult to control, given the multitude of printers and limitless designs. Providing oversight of 3D printed surgical instruments is likely overwhelming, and will require significant adaptation by the Food and Drug Administration. ¹ Second, concerns about sterilizability of printed tools are abound. As previously mentioned, studies have demonstrated that the high printing temperatures associated with FDM can produce a sterile product. ^3,12,13 However, resterilization and the lifespan of such instruments are unknown, and might counteract initial environmental and storage benefits. Lastly, upfront costs for a single printer can be in the hundreds of thousands of dollars, and every printer is slightly different with regard to the printing technique and materials. Therefore, the potential for the need of multiple different printers to cover all possible uses is likely prohibitive currently from a cost standpoint. However, commercial surgical clips can be a significant disposable expense, whereas the actual materials required to 3D print a clip and other items cost mere pennies. Therefore, diverse utilization of 3D printers could make the technology economically advantageous. Further advancement in 3D printers will likely lower costs and perhaps allow for rapid prototyping to have a role in the operating room in the future.

Conclusions

3D printing of surgical clips is currently feasible, although large-scale reproducibility is limited at this time. Despite current limitations with regulation, sterilizability, and cost, printer and material refinement in the future may lead to utilization of printed materials in the operating room.