Fabrication and mechanical testing of polydioxanone hook cross biodegradable self-expandable enteric stent: impact of fabrication density and mechanical properties of the stent
Article information
Abstract
Background/Aims
The mechanical properties of biodegradable stent when fabricated using different number of pins per row of fabrication has been limited. We compared the radial compressive force of polydioxanone (PDO) stent that was fabricated in hook and cross manner, using 13, 17, and 19 pins per row and measure the radial compressive force and ex vivo deployment.
Methods
The PDO stents fabricated by the in-house aluminum mandrel were tested for radial force using plate compression until the stent achieved 50% strain. The relationship between compression force and % strain was calculated. Ex vivo testing of stent expansion against short segment stricture was performed in a pig small intestine compared between PDO hook cross PDO stent and braided metallic stent.
Results
The stent shortening of 16.40%, 31.20% and 19.24% was observed in 13-, 17-, and 19-pin-per-row, respectively. The maximum force to achieve 50% strain were 0.503, 1.168 and 1.008 N for 13, 17, and 19 pins per row, respectively. The stent fabricated using hook and cross pattern demonstrated higher conformability to anatomical stricture when compared with braided stent.
Conclusions
PDO stent fabricated using 17 pins per row demonstrated highest radial force when compared with 13 and 19 pins per row.
INTRODUCTION
Self-expandable stent has become an important tool for the treatment of gastrointestinal stricture.1 In patients with benign conditions, short term stent placement is preferred owing to the high complication rate of long-term stent placement. Most stents that have been used, including metallic stents and plastic stents, provide a good mechanical property but require stent removal. To prevent stent embedment to the surrounding mucosa, covering membrane has been used2 which results in stent migration in up to 20% of patients.3 As a result, biodegradable stent has been adopted for the treatment of benign conditions in which long-lasting effect is not needed. The material that has been used for fabrication of commercially available are biodegradable polymers such as polydioxanone (PDO), polylactide (PLA), and polyglycolide, which are also used as the biodegradable suture. However, the major drawback for these stents is that the mechanical property of these stents is not comparable with the alloys that are used in usual metallic stents. When compared among the biodegradable materials, PDO provided the longest time before degradation.4,5 As a result, PDO is the same material that is used for fabrication of commercially available biodegradable stent.
The commercially available metallic stents are fabricated using various patterns which also affect the mechanical performance of the stent. These can be divided into braided wire or laser-cut type.6 The braided wire includes those in which the wire crosses from one end to the other (braided) and those with complex structure that hooking between wire was intervening with the braiding wire (hook cross type). The latter, so-called hook cross type, results in lower axial force compared to traditional braided catheters while maintaining a similar radial force.7 There has been limited data regarding the mechanical performance of biodegradable stent that fabricated using hook cross design, especially the density of the yarn, that depended on number of pins per row of mandrel that used in fabrication. As a result, we conduct the study aims to evaluate the mechanical properties and deplorability of hand-woven, PDO stent with hook cross pattern when using different number of pins per row of fabrication.
METHODS
Stent fabrication
PDO monofilaments (Meta Biomed Co., Ltd.) with diameter of 0.4 mm and average knot-tensile strength of 39 N is used for fabrication of the stent. The mandrel was made in-house using 23 mm aluminum tube and drilled into 2 mm holes for pin fixation (Fig. 1). The pin number was designed to offer the offset fabrication of the stent, with braid angle of 30° (α=30°) according to the previous study.8 The pins were designed to align parallel to each other. Hook cross pattern of fabrication was selected since this pattern provide lower axial force that enable the stent to conform to the angulated enteric anatomy.9 The stent fabrication pattern is shown in Figure 2 with the calculation of pin numbers and distance between each pin and row height. To enhance the strength of the stent, pattern design that provides 3 cross points was selected (Fig. 3). Figure 4 shows stents were fabricated using mandrel with 13 pin-per-row, 17 pin-per-row, and 19 pin-per-row, each with 3 samples with the same length of 5 cm.
The in-house mandrel for fabrication with attachable 2-mm pins.
Fabrication pattern of polydioxanone hook cross stent.
The scheme of hook and cross fabrication (A) and the final result using polydioxanone (PDO) monofilament (B). Each stent segment are connected by hooking of each row and crossing over in the same roll (crossing part demonstrated in red circle).
Stents fabricated using 13, 17, and 19 pin per roll (A) and a cut hook and cross metallic stent (B).
The difference between fabrication type and the mechanical function of the stent against stricture is shown in Figure 5. Braided stents showed less contact area when compared with stent fabricated using hook cross type. The braided stent functioned as the single unit of radial expansion which is different from the stent fabricated in multiple segments, in which each segment expands independently. Although similar in the fabricated pattern, the 13, 17, and 19 pins per row differs in the height of each cell and the density of the yarn that result in different porosity of the stent.
Difference between fabrication type and the mechanical function of the stent against stricture. Yellow lines mark the contact surface between the stent and stricture and the blue arrow demonstrated the expanding function of the stent.
Process after fabrication
After the manual fabrication, the stent was heat-set at 80 ℃ for 30 minutes to restore the tubular shape after removal from the mandrel.10 Before the usage, the pins were removed, and the stent was removed from the mandrel and stored in a dry container. Stent shortening was observed for each stent.
Mechanical testing
1) Measurement of radial compression
Radial compression is important characteristic of stent as it acts against the enteric stricture. Structural deformation affects the performance when stent is deployed inside the conduits. A radial compression test was conducted on all the stent samples using EZ-S tester, Shimadzu. The radial force was measures using flat plate compression. Three samples of each type of fabrication and the commercially available hook cross metallic stent (Niti S pyloric/duodenal uncovered stent [D-Type]) which was cut in 4 cm were tested. The compression started from resting position to 50% initial diameter with compression rate of 0.2 mm/second for 5 cycles (Fig. 6). The forces used for compression were recorded and the % strain was calculated and compared between each stent. From 5 cycle, the relationship between median force compression and % strain was calculated. As the stent compressed, result in different final length and diameter, % strain (equation 1) was calculated using the following equation. The maximal force that needed to achieve 50% strain was compared between the stent fabricated using 13-, 17-, and 19- pin were compared.
The radial force measurement of polydioxanone stent using EZ-S tester (Shimadzu) (A). The compression was performed until the stent collapsed to 50% of initial diameter (B).
Ex vivo testing
The fabricated PDO stent was inserted to the ex vivo model to evaluate the stent expansion. We selected a pig small intestine for the evaluation of stent expansion as the pig small intestine has small caliber and has a very thin wall that enable visualization though the wall unlike the esophagus or large intestine. The fresh pig small intestine was cut in 30 cm length and fixed at both ends using pins. We used the delivery system for Polyflex stent (Boston Scientific) to deliver the PDO stent into the ex vivo model. After the stent was inserted into the pig’s small intestine, the inner catheter was pushed to deploy the stent and removed. Small intestinal enlargement after the stent insertion.11 Stricture was made using silk tying around the intestine and pulled to create luminal narrowing with 10 mm in diameter. Stent deformation and expansion was observed to evaluate the conformability of the stent to the stricture site by calculating the contact area between the stent and the pig small intestine. The contact area was compared between the 17-, 19-pin, and braided stent at similar length using ImageJ software.
RESULTS
Stent characteristics and stent shortening
Three samples of PDO stents fabricated using 13, 17, and 19-pin per roll were tested. With similar length and diameter, the amount of PDO monofilaments in fabrication of 13, 17, and 19 pins per row stent were 0.47, 0.56, and 0.61 g, respectively.
After removal of the PDO stent from the mandrel, final stent diameter and length were described in Table 1. When compared with initial length of 50 mm, 20% to 30% stent shortening was observed.
Final stent diameter and length
Radial force measurement
The compression study of each PDO stent as shown in Figure 7. When compressed up to 50% of initial diameter, the maximal resistive force of the 13, 17, and 19 pins were 0.503±0.041, 1.168±0.032 and 1.008±0.067 N, respectively. The maximal resistive force of metallic stent was 1.678 N. The result showed highest radial force in metallic stent followed by 17 pins, 19 pins, and 13 pins, respectively. Considering the radial force that needed for the treatment of stricture, 17 and 19 pins were selected for further testing in ex vivo model.
Relationship between radial force and % strain of polydioxanone stent fabricated using 13-, 17-, and 19- pins per roll and the cut commercial metallic stent.
Stent insertion to the pig small intestine
Figure 8 shows stent insertion to the pig small intestine when compared between 17-pin and 19-pin PDO stent when compared to the braided stent (Wallflex Esophageal; Boston Scientific). Immediate stent expansion was observed next to the stricture created by the silk knot. Within the same length, the contact area of the stent was 500.68, 600.85, and 222.84 mm2 in 17-pin, 19-pin, and braided stent, respectively (Fig. 9) with demonstrated gaps between the stent and the surrounding tissue just above the stricture site in the braided stent. These gaps were not observed in the hook cross stent since the stent acts as the corrugated tube that expanded in segmental fashion. These findings reflect the better conformability of the hook cross stent when compared with braid stent.
Stent insertion in the ex vivo model compared between polydioxanone stent (A) and braided metallic stent (B).
The contact area of the stent on the ex vivo model when measured at the same length. Higher contact area was seen in the polydioxanone stent when compared to the braided stent, which showed segmental gap just above the stricture site (arrow).
DISCUSSION
Ideal stent for enteric stricture
The mechanical properties of the stents are the key for the treatment of enteric strictures as it counteracts the stress generated by the surrounding stricture. Apart from the radial expansion, the conformability of the stents to angulation, so-called axial force, is also important as it associated with higher complications.12 The hook cross design is known to conform better to the shape of the esophagus due to their individual sections expanding like a corrugated tube. Despite several studies comparing the mechanical properties of braided stent,8,13 there has been limited study in hook cross type. Recently, various types of polymeric biodegradable stents have been developed for the treatment of gastrointestinal diseases. Although not widely available as a commercial product, these stents showed promising results in the treatment of various conditions.14 Ideally, the desired structure of biodegradable stent should have a mechanical performance comparable to metal stents and biodegradability that can maintain pipe wall stability after implant while raise no concerns for stent migration.
Mechanical properties of the stent
The minimally required radial force that is needed for the expansion of the stricture was unknown and there is no consensus on the optimal force required to maintain esophageal shape after dilation. Moreover, the required mechanical properties, both radial and axial forces, depends on the location and nature of the stricture. The commercially available metallic stent provided 4.25 to 0.66 N/mm at the middle, which was the force that against the stricture site.15 While commercially available biodegradable stents have radial forces of 4 to 5 N and axial forces of 1.8 to 1.9 N.7 Factors associated with radial expansion force included material property, stent design, braiding angle, stent diameter, and the stent porosity.8 With similar material and diameter, the radial force and axial force of metallic stents made from Nitinol or stainless steel ranged from 2.08 to 14.1 N and 0.04 to 0.95 N, respectively, depends on the fabrication pattern.16 The radial force could be calculated using various method such as using the Model TTR2 (Blockwise Engineering LLC)9 or plate compression.17
Fabrication pattern of PDO hook cross stent and the radial force
Our study compared between the fabrication pattern and the radial force measured by plate compression technique. Theoretically, a stent with higher material density (so-called, lower porosity) should exhibit higher radial force. However, our study showed that, despite lower porosity, the 17 pins stent seems to exhibit higher radial force and also a higher shortening rate than the 19 pins stent. We hypothesized that this might be the result of increasing fraction between PDO yarn with the increasing number of crossing PDO yarn, impeding the radial expansion force against external compression. However, our study showed that despite the increase in the yarn density, the radial force was not increased in a linear fashion. We hypothesized that higher yarn density also result in higher friction between the yarn, which impedes the radial expansion.
Although higher radial force theoretically should be more effective for the treatment of stricture, it is associated with higher complication such as chest pain, gastroesophageal reflux, and airway compression.18 Moreover, lower radial force stent was reported to be associated with lower adverse events in the treatment of malignant esophageal stricture after radiation.19 For other polymer stents, the radial compression force of polymer esophageal stent Polyflex was 2.02 to 4.25 N/mm as this model composed of welded polyester embedded in silicone cover.20 In the braided polyester stents, the radial compression force ranges between 1 to 7 N was observed.21 On the other hand, the radial compression force of 3D printed stent using PLA/TPU is in range of 8.2 to 17.7 N at 50% strain.11 Among these stents, only Polyflex is commercially available but became unpopular compared with the fully covered metallic stents. Our study showed that PDO hook cross stent could exhibit up to 1.16 N radial expansion force when compressed to 50% strain, which is within the range of the braided polyester stent in other study.21 Noted that even the hook cross metallic stent with similar length showed maximal radial force of 1.678 N.
Stent shortening and elongation are one of the properties of self-expandable stents22 which varies according to stent design. In commercially available esophageal stent, stent shortening was reported ranges from 0% to 10% in Z type stent to 30% to 40% in braided stent.23 Our study showed 16.4% to 31.2% stent shortening which is similar to what observed in commercial stent. To eliminate the effect of stent shortening, we calculated the change in compressive force against % strain instead of compressive diameter, which might be affected by stent deformation. The stent shortening of 17 pins was higher than that from 13 and 19 pins, which might imply the higher recoiled of shorter PDO yarn in each segment of fabrication when compared to 13 pins and less friction between the crossing yarn when compared to 19 pins. The amount of PDO yarn in 19 pins stents was 0.61 gram compared to 0.47 gram in 13 pin stent, which reflects the higher density of the yarn when the stent is fabricated using a higher number of pins.
Stent expansion in ex vivo model
In the ex vivo model, the hook cross stent showed the better conformation of the stent when compared to the metallic braided stent. When a short narrowing of the lumen is created, it changes the shape into dumbbell shape, which is similar to the lumen-apposing metal stent (LAMS) that has become another choice of treatment especially in short segment stricture.24 Despite the benefit of better patient tolerance, LAMS placement still associated with a high migration rate of up to 40%.25 The conformability of the stent could be demonstrated by measuring the apposition area between the stent and the surrounding structure.24 By using a similar method, our study demonstrated that contact area between the stent and the pig intestine was higher in hook cross stent, which reflect the better conformation to the stricture. With the benefit of biodegradability, the PDO hook cross stent might be one of the future choices for the treatment of short segment enteric stricture.
Our study demonstrated that, although higher yarn density should increase the strength of the fabricated stent, this may not be observed in a linear fashion as the increasing yarn density would result in higher friction. This technique also results in stents that have a lower radial force that was reported to correlated with lower complication.
However, our study still has several limitations. Firstly, the PDO stents were fabricated manually, which might result in non-uniform tension of the yarn during fabrication. Secondly, the stent fabrication was limited to the odd number, which might not be able to demonstrate the exact correlation between pin number and radial force. Also, the stent that was tested for fabrication was shorter than commercially available stent, which might result in lower radial force than expected. With longer stent fabrication, the radial force of the stent might be higher than that demonstrated in this study.
In conclusion, our study demonstrated the correlation between the number of pin-per-row and the radial force. Although a higher number of yarn density should correlate with higher radial force, a higher number of yarns might cause higher friction which impact the radial force of the stent. When compared between the 13, 17, and 19 pin hook cross stent, 17 pin stent provide the highest radial force while maintained good conformability when compared with braided stent.
Notes
Conflicts of Interest
The authors have no potential conflicts of interest.
Funding
This work was supported by Faculty of Engineering, Chulalongkorn University, Bangkok, Thailand.
Acknowledgments
The authors would like to thanks Ms. Sarinthorn Boonkruephan for her assistance in laboratory testing.
Declaration of generative AI and AI-assisted technologies in the writing process. During the preparation of this work the author(s) used Gemini AI in order to improve the readability and language polishing. After using this tool/service, the author(s) reviewed and edited the content as needed and take(s) full responsibility for the content of the published article.
Author Contributions
Conceptualization: all authors; Data curation: TC; Formal analysis: TC, SK; Investigation: TC; Methodology: all authors; Writing–original draft: all authors; Writing–review & editing: all authors.