Flexible, Stretchable and Implantable PDMS ... - Springer Link

Biomed Eng Lett (2011) 1:199-203 DOI 10.1007/s13534-011-0033-8

ORIGINAL ARTICLE

Flexible, Stretchable and Implantable PDMS Encapsulated Cable for Implantable Medical Device Sung Hwan Kim, Jin-Hee Moon, Jeong Hun Kim, Sung Min Jeong and Sang-Hoon Lee

Received: 9 August 2011 / Revised: 26 August 2011 / Accepted: 29 August 2011 © The Korean Society of Medical & Biological Engineering and Springer 2011

Abstract Purpose Diverse commercial implantable medical devices were developed for the convenience and life-quality of patients, and tether-free diagnostics and therapeutics. Those devices need implantable cable, which connects each part of their devices for transcutaneous energy or signal transfer. For prolonged implantation into human body, it should be safe, robust, and be long-term operable without failure. In this paper, we introduce an implantable PDMS-coated cable. Methods By using PDMS as an encapsulation material, we developed a biocompatible, flexible and durable cable. Several tests were carried out for the evaluation of cable performance. Leakage test confirmed that coating using PDMS was sufficient to prevent the invasion of body fluid to conducting wire. Tensile test, torsion-durability test and bending-durability test demonstrated its mechanical durability and robustness to motion. The resistance variation recorded in the durability test was monitored to verify the electrical stability during movement of cable. Experiments of subcutaneous tissue implantation for 8 weeks were performed to observe the degree of biocompatibility as an implantable cable. Results & conclusions Our cable was mechanically stable and biocompatible enough to be used for long term implantable medical devices. Keywords Implantable cable, PDMS, Biocompatibility, Electrical connection

Sung Hwan Kim†, Jin-Hee Moon†, Jeong Hun Kim, Sung Min Jeong, Sang-Hoon Lee ( ) Department of Biomedical Engineering, College of Health Sciences, Korea University, Seoul 136-100, Korea Tel : +82-2-940-2881 / Fax : +82-2-921-6818 E-mail : [email protected] †

equally contributed author

INTRODUCTION For the patients’ convenience, life-quality and for tether-free diagnostics and therapeutics, the implantable medical device (IMD) has been highly focused and diverse commercial IMDs were developed. To implant IMDs in human body without complications, it has several critical requirements such as autoclavability, nontoxicity, biocompatibility and flexibility to shape freely to the change of patient’s posture [1]. Furthermore, the IMDs should be safe, robust, and long-term operable without failure. Recent progress of fabrication and packaging technology enhanced the durability, mechanical stability and robustness of IMDs. However, for the enhancement both of transcutaneous energy or signal transfer and of the accessibility to the target position for sensing or stimulation, complicated IMDs are generally required to be divided into two or more parts. The separation into some parts enhances patients’ activity and solves the space shortage of targeted organs or tissues when all-in-one system is implanted in. However, the separation of devices inevitably needs implantable cables to connect among parts [2, 3]. The cables are generally composed of core conductors made of several metal alloys and insulation material resembling a ceramic, carbon or polymer barrier to prevent it from direct contact with body fluids [1, 4]. Therefore, the safety of implantable cables is a critical factor of IMDs for the failurefree operation. However, implantable cable has some significant problems including degradation of insulation material by metal-ion oxidation and lack of fatigue resistance [4-8]. In case of transvenous defibrillation lead, many problems were reported accordingly [5, 9-11]. Long-term implantation for average 934 days caused 15% failure, and the major lead complications were insulation defects (56%) and lead fractures (12%) [10]. To address this limit, complex materials of polyurethane (PU) and other materials were used as an insulation material for commercial implantable lead of defibrillator. Despite of its several advantages, PU is

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not as stretchable and degradable [6, 7, 12-17]. Stretchable, biocompatible, resistant to body fluid, flexible and softer cables are needed for small IMDs to be implanted into limited space surrounded by vulnerable tissue. In this paper, we developed the implantable cable insulated with polydimethylsiloxane (PDMS) to overcome the problem of conventional implantable cables. PDMS is one of the most widely used silicon-based organic polymers, and broadly applied in diverse biomedical fields including contact lens and microfabricated chip. PDMS is transparent, a chemically inert material, possessing high biocompatibility, easy fabricability, high elasticity and softness [18]. The copper wire was wound spirally and encapsulated with PDMS. The high flexibility and biocompatibility of PDMS can increase the robustness to external forces and foreign-body sensation, and we investigated such properties through several experiments. The mechanical and electrical properties were evaluated by measuring electrical resistance and durability to mechanical deformation and contact impedance. The biocompatibility was tested by implanting cables in the subcutaneous tissue of mouse for 8 weeks and by analyzing the surrounding tissue through the hematoxylin and eosin (H&E) staining.

MATERIALS AND METHODS Fabrication of implantable cable encapsulated with PDMS The fabrication of cable was processed in a total of six steps. (Fig. 1): Firstly, the conducting wire (UL-1332 FEPW, fluororesin insulated 100% copper wire, Totoku Co., Japan; radius 650 µm) was wound on an core cylindrical iron rod (length: 15 cm, radius 0.6 mm) (Fig. 1a). Winding pitch was determined to be 1.8 mm considering the stretching capability

Fig. 1. Process of fabrication. (a) Winding of wire around cylindrical core. (b) Insertion of core wound wire into tube mold. (c) Removal of core from tube mold. (d) Injection of PDMS solution into tube mold. (e) Baking on hotplate at 90oC during 90 minutes. (f) Separation of completed cable encapsulated with PDMS from tube mold.

Biomed Eng Lett (2011) 1:199-203

of PDMS and electrical resistance. Secondly, the iron rod wound by the wire was covered with tube mold (TP-030, IXAK®; length 15 cm, inner radius 0.85mm) (Fig. 1b). Then, the iron rod was removed from the tube mold (Fig. 1c). To maintain steady winding pitch, one end of the conducting wire was slightly fixed, and the rod was carefully pulled out. Thereafter, the encapsulation of the wound wire with PDMS (Dow Corning sylgard® 184) was performed. Syringe pump (KDS100 Syringe Pump, KD Scientific Co., US) was used to inject degased PDMS solution (10:1 weight ratio mixture of PDMS and curing agent) into the tube at flow rate of 11.3 µL/s (Fig. 1d). When the tube mold was filled up with PDMS precursor, it was baked on hot plate (Digital hot plate HP-1S, Labotec co., Korea) at 90oC during 90 minutes (Fig. 1e). After baking, the spiral cable encapsulated with PDMS was taken out from the tube mold (Fig. 1f). Leakage test The leakage test was conducted to prove that encapsulation of PDMS can prevent leakage current in the body. Experiment was conducted for two groups; uncoated wire and cable encapsulated with PDMS. The sample was dipped into NaCl solution (0.9%), and an iron rod was used as a cathode. Multi-meter (34401A Digit multi-meter, Agilent Co., USA) was used to measure the resistance between cathode and the sample. If encapsulated wire is exposed to the solution, the leakage current may start to flow, in turn causing resistance change. Tensile test Tensile test is one of general methods to characterize the mechanical properties of cable to the motion caused by muscle contraction and external forces. Implantable cables encapsulated with PDMS were prepared. For the comparison study, same sized and shaped PDMS cable without copper wire were prepared. For each groups, five samples were prepared. One end of each sample was attached to sensor of digital force meter (FGN-5B, NIDEC-SHIMPO Co., Japan) and the other end was dragged by translational stage until the sample was broken. Force was measured at every 0.5 cm extension, and Young’s modulus and the yield stress were calculated. Also, the resistance change was measured at every 1 cm extension to monitor the disconnection of wire during tensile test. Torsion test Torsion-stress fatigue test was conducted for a week by monitoring the variation of resistance in every two days. Two-phase step motor (PK233PA, INA Oriental Motor Co., Korea) was programmed to rotate 360 degrees for 2 seconds in a clockwise motion and counter-clockwise alternatively. One end of the sample was attached to spindle of motor,


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while the opposite end was fixed. Test was performed for 7 days, by twisting the samples over 151,200 times (7days × 24 hours × 60 minutes × 60 seconds / 4 seconds = 151,200). After test, leakage current was measured to detect the damage of the cable. The resistance variation to the degree of twisting was measured. The cable was twisted from 180 degrees to 1,820 degrees (180 × 9). At every 180 degrees turn, resistance was measured. Bending durability test Bending durability test was conducted to check fatigue resistance over consecutive bend force for a week, and the resistance variation was also measured in every 2 days. We established home-made durability test equipment using twophase step. When the motor rotates, the samples were bent at angle of 90 degrees per second. The estimated total number of bending was over 604,800 times (7 days × 24 hours × 60 minutes × 60 seconds = 604,800). After test, leakage current was measured to detect the damage of the cable. Subcutaneous implanting test All procedures were performed in accordance to protocols approved by the Korea University Institutional Animal Care &Use Committee (IACUC). To verify the biocompatibility of the cable, implanting test was performed for 8 weeks. Implantable cable was cut to the length of 2 mm and implanted underneath of epidermis. Institute of Cancer Research (ICR) mice (3 weeks aged) was prepared and anesthetized by diethyl ether; the prepared samples were implanted at the two regions beneath back skin. After 1, 2, 4, 8 weeks, tissues of dimensions 5 × 5 × 5 mm were retrieved. To examine the morphology of isolated sample, the tissues were fixed in 10% neutral and isotonic buffered formalin for 12 hours and the samples were paraffinfixed after being dehydrated. The samples were cut into 0.8mm-thick slices and the sections were stained with H&E for the observation of inflammatory response.

Fig. 2. The completed cable encapsulated by PDMS. The pitch and thickness were almost constant, and the length was 15 cm (mold’s length).

Leakage test was conducted using 10 samples (5 samples were conductors without PDMS coating and 5 samples were implantable cable encapsulated by PDMS). The conducting wire was coated with thin enamel. Although enamel coating was applied, the resistance between wire and saline solution was measured to about 1~9 MΩ, which involves leakage current generated in the non-encapsulated cable. However, in the PDMS encapsulated implantable cable, we did not observe any leakage current indicating that the PDMS protects the wire completely from body fluid. From the tensile test, strain-stress curve was plotted as shown in Fig. 3. The graph demonstrates that the implantable cable could be elongated ~ 140% of its original size, and ultimate tensile strength (UTS) was 488.1 kPa. In the case of same sized PDMS fiber without conducting wire, it could stretch more than 170% of its own size. Young’s modulus of PDMS fiber was 0.612 MPa, while that of the implantable

RESULTS AND DISCUSSIONS The extendable, flexible and implantable cable was successfully fabricated as intended (Fig. 2). Its length was 15 cm, radius 0.8 mm and resistance was approximately 0.09 Ω/cm. The resistance would be affected by the pitch of winding because short pitch could cause increase length of wound wire. The size is sufficient to connect the separated IMDs in spite of narrow intra-body space surrounded by soft tissue. PDMS encapsulated the spiral conductor completely and cracks were not observed by the microscopic observance. Fabricated cable could be stably connected to any implantable device, and all process was completed within 2 hours.

Fig. 3. Stress-strain curve of the cable coated by PDMS and PDMS pole without wound wire (control). For 5 samples per each case, stress(F/A) at each strain(∆L/L) was measured. Estimated line was calculated using least square method (UTS : ultimate tensile strength).

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Fig. 4. Resistance variation of each mechanical test. (a) Resistance variation by strain. For 5 samples, resistance at each strain(∆L/L) was measured. (b) Resistance change by bending-durability test. Resistance was measured at every 2 days. (c) Resistance variation by twisting the wire. The wire was twisted from 180 degrees to 1,620 degrees (180 × 9). At every 200 degrees angle, resistance was measured. (d) Resistance change by torsion-durability test. Resistance was measured at every 2 days.

cable with conducting wire was 1.15 MPa. The reason could be that the conducting wire is tightly restricted by surrounding PDMS, which may prevent the high extension of wire. The resistance of encapsulated wire was not changed by elongation (Fig. 4a). After conducting the stretching test, the current was measured to monitor the leakage current, and no significant changes were observed, implying that the wire was not damaged by the stretching. Bending durability test was carried out for a week, using 5 samples. Resistance in every 2 days did not change during experiments (Fig. 4b). Torsion-resistance test showed that electrical conductivity of the cable was not affected by twisting (Fig. 4c). Surprisingly,

there was no breakage of cable even though it was twisted for 20 turns (72,000 degrees). Such superior property of torsion resistance may be from the spiral structure of conducting wire, indicating extensive industrial applications requiring high torsion. Torsion-durability test was conducted for 7 days using 5 samples. In every 2 days, resistance was measured, and it was almost constant (Fig. 4d). After torsion test, encapsulating PDMS showed no visible damages and the conductance of cable was maintained consistently. The leakage test proved the strong protection capability of implantable wire from body fluid. Visible damage of encapsulating PDMS was not observed, and the conductance

Fig. 5. Microscopic images of implanted wire and surrounded hypodermic tissue of mouse 4 weeks after implantation. P regions (white dotted line) are the empty space where cables coated with PDMS were located. The fibrosis (white arrow) are shown 2, 4, 8 weeks after implantation (a, b, c are pictures after 2, 4, 8 weeks from implantations beneath epidermis, respectively).


of wire was maintained well during test. The leakage test confirmed encapsulation state was enough to keep it safe from current leakage. Generally, PDMS has low fatigue resistance and mechanical strength than PU or other materials. Despite of these properties, the proposed PDMS wire was relatively durable to large bending and torsion force. In addition, the electrical disconnection was not observed, possibly due to the spiral winding of conducting wire. For subcutaneous implants test, tissues of 1, 2, 4, 8 weeks after implantation were examined. As expected from several reports [19, 20], any inflammation was not observed in all tissue samples. For the detail inspection, all samples were H&E-stained and observed through the microscopy. During the first week after implantation, fibrosis and inflammation were not observed. At 2, 4, 8 weeks after implantation, fibrosis was observed (see arrows at Fig. 5) but inflammation was not observed. This result indicates that PDMS do not cause any inflammatory trouble to surrounding tissue.

CONCLUSION In this paper, implantable cable encapsulated with PDMS was successfully developed and its quantitative properties for implantable use were examined. PDMS encapsulation was proved to be a protective layer to exposure of body fluids. Mechanical durability test demonstrated the endurance of cable from the continuous muscle movement. The flexibility and softness offer advantages to use it as an implantable device in the movement of tissues. In vivo tests for 8 weeks also showed the biocompatibility of PDMS materials. Due to the many advantages, the proposed technology has several industrial applications including robotics, movable machines, mechanical artificial prosthesis and automation in biomedical research.

ACKNOWLEDGEMENT This study was supported by a grant of the Korea Healthcare technology R&D Project, Ministry for Health, Welfare and Family Affairs, Republic of Korea (A092106).

REFERENCES [1] Donaldson PE. The Cooper cable: an implantable multiconductor cable for neurological prostheses. Med Biol Eng Comput. 1983; 21(3):371-4. [2] Marino JC. Flexible connector for implantable wiring harness. In.: US Patent App. 20,080/009,198; 2005. [3] Kaplan PM, Cross Jr TE, Bolea SL, Olsen JM. Implantable medical lead with axially oriented coiled wire conductors. In.:

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Google Patents; 2010. [4] O’brien RC. Implantable electrical lead wire. In.: EP Patent 1,547,647; 2005. [5] Kitamura S, Satomi K, Kurita T, Shimizu W, Suyama K, Aihara N, Niwaya K, Kobayashi J, Kamakura S. Long-term follow-up of transvenous defibrillation leads: high incidence of fracture in coaxial polyurethane lead. Circ J. 2006; 70(3):273-7. [6] Ward B, Anderson J, Ebert M, McVenes R, Stokes K. In vivo biostability of polysiloxane polyether polyurethanes: resistance to metal ion oxidation. J Biomed Mater Res A. 2006; 77(2):380-9. [7] Ward R, Anderson J, McVenes R, Stokes K. In vivo biostability of polyether polyurethanes with fluoropolymer and polyethylene oxide surface modifying endgroups; resistance to metal ion oxidation. J Biomed Mater Res A. 2007; 80(1):34-44. [8] Szycher M, Reed AM, Siciliano AA. In vivo testing of a biostable polyurethane. J Biomat Appl. 1991; 6(2):110-30. [9] Ellenbogen KA, Wood MA, Shepard RK, Clemo HF, Vaughn T, Holloman K, Dow M, Leffler J, Abeyratne A, Verness D. Detection and management of an implantable cardioverter defibrillator lead failure: incidence and clinical implications. J Am Coll Cardiol. 2003; 41(1):73-80. [10] Kleemann T, Becker T, Doenges K, Vater M, Senges J, Schneider S, Saggau W, Weisse U, Seidl K. Annual rate of transvenous defibrillation lead defects in implantable cardioverterdefibrillators over a period of >10 years. Circ. 2007; 115(19): 2474-80. [11] Hauser RG, Cannom D, Hayes DL, Parsonnet V, Hayes J, Ratliff N, 3rd, Tyers GF, Epstein AE, Vlay SC, Furman S et al. Longterm structural failure of coaxial polyurethane implantable cardioverter defibrillator leads. Pacing Clin Electrophysiol. 2002; 25(6):879-82. [12] Luo YL, Nan YF, Xu F, Chen YS, Zhao P. Degradation behavior and biocompatibility of PEG/PANI-derived polyurethane copolymers. J Biomat Sci-Polym E. 2010; 21(8-9):1143-72. [13] Stachelek SJ, Alferiev I, Ueda M, Eckels EC, Gleason KT, Levy RJ. Prevention of polyurethane oxidative degradation with phenolic antioxidants covalently attached to the hard segments: structure-function relationships. J Biomed Mater Res A. 2010; 94(3):751-9. [14] Zuidema J, van Minnen B, Span MM, Hissink CE, van Kooten TG, Bos RR. In vitro degradation of a biodegradable polyurethane foam, based on 1,4-butanediisocyanate: a three-year study at physiological and elevated temperature. J Biomed Mater Res A. 2009; 90(3):920-30. [15] Stachelek SJ, Alferiev I, Choi H, Chan CW, Zubiate B, Sacks M, Composto R, Chen IW, Levy RJ. Prevention of oxidative degradation of polyurethane by covalent attachment of di-tertbutylphenol residues. J Biomed Mater Res A. 2006; 78(4):653-61. [16] van Minnen B, Stegenga B, van Leeuwen MB, van Kooten TG, Bos RR. A long-term in vitro biocompatibility study of a biodegradable polyurethane and its degradation products. J Biomed Mater Res A. 2006; 76(2):377-85. [17] Gorna K, Gogolewski S. Preparation, degradation, and calcification of biodegradable polyurethane foams for bone graft substitutes. J Biomed Mater Res A. 2003; 67(3):813-27. [18] Belanger MC, Marois Y. Hemocompatibility, biocompatibility, inflammatory and in vivo studies of primary reference materials low-density polyethylene and polydimethylsiloxane: a review. J Biomed Mater Res. 2001; 58(5):467-77. [19] Park J, Yoo S, Lee E-J, Lee D, Kim J, Lee S-H. Increased poly(dimethylsiloxane) stiffness improves viability and morphology of mouse fibroblast cells. BioChip J. 2010; 4(3):230-6. [20] Peterson SL, McDonald A, Gourley PL, Sasaki DY. Poly(dimethylsiloxane) thin films as biocompatible coatings for microfluidic devices: Cell culture and flow studies with glial cells. J Biomed Mater Res Part A. 2005; 72A(1):10-8.