End-stage renal disease affects more than 425,000 Americans and continues to increase in prevalence at 8% annually.1,2 Renal transplantation (the gold standard) is hindered by a shortage of donor organs, whereas dialysis is expensive, inconvenient, and confers significant morbidity and mortality. Miniaturized, implantable, or wearable renal replacement systems could improve the implementation of newer dialysis prescriptions, such as prolonged and daily therapies at home.3–5 Tissue engineering may provide the solution for vascular filtration interface issues.6
Our group has focused on developing a functional hemofiltration membrane as a part of perfecting the miniaturization of renal replacement therapy.7 Wide variation in pore size within conventional polymer membranes leads to imprecise filtration characteristics. As the mean pore size is reduced, solutes above the desired molecular weight cutoff of the membrane are effectively retained, but the hydraulic permeability of the membrane is reduced.8,9 Engineering narrower pore size distribution maximizes the mean pore size of the membrane and allows for sharper transitions from passage to retention.7–12
Microelectromechanical Systems (MEMS) technology-associated silicon micromachining techniques provide an attractive toolkit for the development of miniaturized implantable medical devices. MEMS technology offers nanometer-scale control of device features using an industrially mature, high-volume, low-unit cost manufacturing process.13 Our group developed a novel ultrafiltration membrane constructed from silicon and related MEMS materials. This next generation hemofilter has well-controlled pore sizes and shapes to improve hydraulic permeability and membrane selectivity, and it will be miniaturized for use in wearable or implantable renal replacement.7
A potential significant downside to the use of silicon as a hemofiltration membrane is the native layer of silicon dioxide (silica, SiO2) that forms spontaneously under atmospheric conditions. Silica is used to stimulate clotting in clinical assays, and blood contact with the hydroxyl groups of oxidized silicon has the potential to initiate blood coagulation by contact activation. Another coagulation pathway involves protein adsorption and conformation change on artificial surfaces, which can initiate clot formation by promoting platelet attachment and degranulation.14,15 In vitro experiments have shown that platelet adhesion on bare silicon and its derivatives was significantly greater than platelet adhesion on polyurethane.16 Despite the advantages of MEMS technology in manufacturing miniaturized devices, chemical or biological surface modification of surface oxides may be necessary for deployment of silicon-based medical devices in the blood. Nevertheless, the native oxide does have certain advantages as an initial coating. First, its formation is spontaneous on exposure of silicon to air and self-limiting in thickness to 1–2 nm. Consequently, the need for sophisticated furnaces, which is used for growth of thicker films, is eliminated. In contrast, the thickness of Si3N4, which has also been used to coat MEMS devices, is typically in the 100–2000 nm range and requires specialized vapor deposition equipment. Second, the oxide can readily provide a high density of silanol (Si-OH) groups for chemical modification of the silicon surface via self-assembly techniques.
Surface modification of silicon with polyethylene glycol (PEG) films has been attempted to reduce protein adsorption.17–20 Although PEG-coated silicon has shown evidence of biocompatibility, much of this evidence centers around soft tissue implants, and intravascular placement has not been specifically evaluated.21,22
We developed an in vivo technique to assess the thrombosis and inflammation characteristics of materials that may be used in the construction of intravascular or transvascular implants. Small, thin samples of implant materials are inserted transversely through rat femoral veins, so that different segments of the implant are simultaneously exposed to adventitial, soft tissue, and blood environments. The implants and vein segments are retrieved for analysis after days to months in vivo. This permits simultaneous assessment of tissue-material interactions in three tissue types without threatening the survival of the animal should clot or inflammation occur. We tested bare and surface-modified silicon implants using this assay.
Materials and Methods
Approval for animal use was granted by the University Committee for the Use and Care of Animals in accordance with the Guide for the Care and Use of Laboratory Animals (NIH publication 86-23, 1986).
Preparation of Silicon Implants
The silicon implants were a gift of H-Cubed Inc. (Olmsted Falls, OH), who performed the micromachining of silicon wafers into miniature “darts.” These implants consisted of a tapered shaft 500-μm wide at the base, 60-μm thick, and 2–4 mm in length. The silicon surface of some implants was covalently modified with PEG using a previously reported solution phase method, modified to omit all sonication steps, and continued the PEG deposition for 12 hours.23 The technique for PEG attachment was a single-step mechanism to covalently couple silicon surface silanol groups (Si-OH) to a PEG polymer through a trimethoxysilane group, thereby forming a Si-O-Si-PEG sequence by a methanol dehydration reaction. PEG coatings added <5 nm to the overall thickness as determined by ellipsometry. All surfaces were generally hydrophilic with water contact angles 10°–20° and approximately 35° on bare and PEG-coated darts, respectively. Five silicon darts were surface modified with PEG, whereas six had no coating applied (termed “bare silicon”).
Eleven silicon darts and 11 controls were implanted in adult, male Fisher F344 rats weighing approximately 300 g. Rats were anesthetized by intraperitoneal injection of ketamine (75 mg/kg, Sigma-Aldrich, St. Louis, MI). The inguinal area was shaved, prepped, and draped sterilely. A 2-cm longitudinal incision was made overlying the femoral vessels, and the femoral artery and vein were exposed using an operating microscope. A dart was then inserted through the diameter of the femoral vein transversely, such that blood flow was not restricted (Figure 1A). One implant was placed per groin. After transvascular placement, the implants were loosely covered with a portion of the inguinal fat, and the incision was closed in two layers. All implants remained in vivo for 4 weeks. Animals were allowed to eat and drink ad libitum.
Gore-Tex Sutures and Stainless Steel Controls
Polytetrafluoroethylene (Gore-Tex) sutures (CV-6, TTc-9 taper needle; Gore-Tex, Inc., Elkton, MD) and stainless steel needles (C-1 taper needle; Ethicon Inc., Cornelia, GA) were implanted as controls. Stainless steel and Gore-Tex were selected as control materials, because each is commonly used in medical implants and is considered to induce minimal tissue inflammation. Each Gore-Tex suture was placed through the diameter of the femoral vein similar to the silicon dart implantation. After the suture was placed, the needle was removed, and the suture was trimmed to leave lengths of suture protruding from each side of the vein (Figure 1B). Each stainless steel needle was placed through the vein, and the sharp needle tip was cut off, leaving the body of the needle in place through the vein wall (Figure 1C).
Explantation of Implants
After 4 weeks in vivo, all implants and their associated vein segments were explanted. The points at which each implant traversed the vein were marked for later identification, and the implants were removed from the veins.
Histologic Analysis of Femoral Vein and Surrounding Tissues
Femoral vein segments were placed into a 4% paraformaldehyde solution for 2 hours and then into 70% ethanol until embedding. Each fixed femoral vein was paraffin embedded, serially cross-sectioned in 5-μm sections, and stained with hematoxylin and eosin. The degree of thrombus within the vein lumen and perivascular inflammation was assessed by a single, blinded viewer. Each vein was assigned subjective grades based on the overall degree of thrombus and tissue inflammation seen (thrombus: “—,” none; “+,” small [<25% of vein lumen]; “++,” moderate [25%–50% of vein lumen]; “+++,” marked [50%–100% of vein lumen] and inflammation: “—,” none; “+,” mild; “++,” moderate; “+++,” marked).
Analysis of Silicon Implants
Each silicon implant was explanted, gently rinsed with phosphate-buffered saline, incubated in 4% paraformaldehyde for 1 hour, and then held for processing in 70% ethanol. Each explant was examined by scanning electron microscopy (SEM). As our study was intended to provide a preliminary assessment of silicon biocompatibility, the small sample sizes and qualitative nature of the observations are not suited to statistical significance testing.
All rats survived to explantation. On exploration, some implants were found to have migrated out of their vein lumens. Such veins were not evaluated. Fully evaluated implants and vein segments were as follows: bare silicon, n = 4; PEG-coated silicon, n = 4; Gore-Tex, n = 2; stainless steel, n = 3. At explantation, all veins were found to be patent by strip test. On gross examination, there was no inflammation, and very few fibrous adhesions were seen surrounding the implants (Figure 1,D–F).
The points of entry of all implants were determined histologically by identifying the mild disturbances in the vein walls and by the ink markings placed during explantation. Cross-sections through the veins at these locations were chosen for grading.
One of four vein segments associated with a bare silicon implant developed a small thrombus (<5% of cross-sectional area). None of the four vein segments implanted with PEG-coated silicon developed thrombi (Table 1; Figure 2, A and B). All three vein segments associated with the stainless steel controls developed moderate (25%–50% of cross-sectional area) thrombi (Table 1; Figure 2C). The two vein segments implanted with the Gore-Tex controls also developed moderate thrombi (Table 1; Figure 2D).
There was no inflammation seen around any of the four vein segments associated with the bare silicon implants (Figure 2A). A mild perivascular inflammatory response was noted around each of the four vein segments associated with the PEG-coated silicon implants (Table 1 and Figure 2B). By contrast, all the three veins implanted with the stainless steel controls developed moderate perivascular inflammation, and each of the two veins associated with the Gore-Tex controls developed marked inflammation (Table 1 and Figure 2, C and D). The inflammatory response was focal—concentrated in the soft tissue immediately surrounding the vein wall near the point of implant entry. An examination revealed a recruitment of macrophages and lymphocytes to the soft tissues surrounding the control materials.
Scanning Electron Microscopy
Surface analysis of the silicon implants by SEM revealed that the bare silicon material seemed to have significant adherent thrombi on the smooth surfaces of the intravascular portions of the implants (Figure 3, A and C), characterized by the red blood cells, platelets, and fibrous debris. In contrast, PEG-treated silicon implants showed occasional adherent platelets and red blood cells, but no evidence of organizing platelet-red cell-fibrin thrombi (Figure 3, B and D). We did not perform specific tests on explanted darts beyond gross inspection and SEM examination to investigate changes in physicochemical properties. There was no observable degradation in the thickness of the explanted silicon darts. The 28-day stability of the PEG coating was excellent as evidenced by the absence of adherent thrombi and the presence of only a few platelet and red blood cells on the darts.
Similarly, silicon implants that had migrated outside of the vessels (and extravascular segments of those that remained in place) had adherent material (data not shown).
In the design of an implantable bioartificial kidney, silicon nanoporous hemofilters are desirable because of their narrow pore size range and enhanced membrane selectivity and permeability.24 However, thrombosis continues to plague conventional hemodialysis patients and will threaten any implantable hemofilter. A significant disadvantage of silicon is its potential to initiate the clotting cascade.
In this study, silicon showed excellent blood and soft tissue biocompatibility. Only one small thrombus was witnessed within a vein associated with a bare silicon implant. It is possible that without PEG coating, the bare silicon induced the thrombus. A confounding variable is the minor intimal damage caused during implantation, an inherent limitation of the transluminal model. PEG is well known for its properties of low protein adsorption and cell adhesion. The ordered arrangement of water molecules surrounding each PEG chain provides a hydrated shell that limits the protein adsorption to the silicon surface, and thus, recognition by the platelets.
Electron microscopic surface analysis revealed that PEG surface modification significantly reduced the degree of thrombosis on the intraluminal silicon surfaces. This finding supports the theory that the thrombus seen in the vein associated with the bare silicon implant developed by extension from the surface of the untreated silicon. In contrast, moderate thrombi were seen in every vein in which a control (Gore-Tex suture and stainless steel) was implanted. The results of this study provide reasonable evidence that silicon, as a structural material for nanoporous membranes in an implantable hemofilter, is less thrombogenic than other common vascularly implanted materials.
Induction of inflammation is another potential undesirable effect of an implant material and can lead to the eventual failure or need for explantation of an implanted device. Other studies have demonstrated the biocompatibility of silicon for use in implantable medical devices.25–28 It was demonstrated that repeated electrical activation of an in vivo MEMS drug delivery device did not induce a significant inflammatory response.29 Furthermore, the leukocytes were rarely present with long-term, in vivo silicon implants.30 In this study, both bare silicon and PEG-coated implants induced less perivascular tissue inflammation than did either of the control materials. A marked inflammatory reaction was observed in the tissues surrounding Gore-Tex sutures, and a moderate response was seen in the perivascular tissues around stainless steel controls.
The very mild proinflammatory effect of PEG-coated darts is likely because of free PEG moieties, which resulted from the hydrolysis of the silane linkers attached to the silicon surface.
The in vivo model used in this study allowed us to collect simultaneous data on the interactions of implants with the blood, vein wall, and perivascular soft tissues. A key advantage of this model is that the blood contact characteristics of any implant material of this relative size can be assessed by direct transluminal implantation, while allowing for continued vascular patency. Many conventional methods of evaluating tissue-material interactions of implants rely on in vitro assays to examine thrombosis and inflammation. Although valuable information can be gained to provide preliminary guidance on implant material choices, the in vitro assays do not fully capture the interactions between the in vivo milieu and the implant material. Traditionally, long duration in vivo biocompatibility investigations require use of large animals and placement of the implant materials in customized packages. In contrast, the method presented in the article is generic to materials that can be machined into a thin slivers or dart shapes and implemented in rodents. The insertion of the material through the vascular wall and flowing blood allows simultaneous testing of material-tissue interactions at two sites. Therefore, our method serves as a relatively low-cost intermediate step between in vitro assays and large animal biocompatibility investigations. A technical challenge of this study was that a number of samples migrated out of their original position through the vein lumens into the subcutaneous tissues. Neither the Gore-Tex nor the stainless steel controls had been fixed to the surrounding tissues on placement through the veins. Technical refinements, such as tying the ends of the suture together, and bending the needles into a circular configuration should be considered in future studies using this transluminal model.
Bare and PEG-coated silicon implants provoked minimal thrombosis and no tissue inflammation compared with control materials commonly used in vascular applications. In addition, the PEG coating provided resistance to biofouling of the silicon surfaces in contact with active blood flow. The transluminal model presented herein proved to be a useful tool to evaluate the biocompatibility and tissue-material interactions of implantable materials. These results are promising evidence for the use of silicon nanopore membranes in the development of an implantable hemofilter. The eventual realization of a bioartificial kidney could lead to a bridge-to-transplant or destination therapy for end-stage renal disease.
The authors thank the following individuals for their contributions to this work: Dr. Wen Zhang, M.D. provided excellent technical assistance during the implantation/explantation procedures; Anna Dubnisheva assisted with the surface analysis of the silicon implants; Gore-Tex Inc. provided the Gore-Tex suture; and Kenneth Goldman of H-Cubed Inc. provided the silicon implants.
Supported by NIH Grant R01EB008049 by the National Institute of Biomedical Imaging and Bioengineering.
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