The Durability of Silicone versus Latex Mock Arteries

Biomedical Science Instrumentation, 37, pp. 305-312, (2001)

CONTI, JC1, Strope, ER, Goldenberg, LM, Price, KS

Dynatek Dalta Scientific Instruments, 105 E. Fourth St., Galena, MO 65656

1 Department of Physics, Southwest Missouri State University

Abstract

Latex mock arteries used in medical device testing allow researchers to evaluate mechanical characteristics of intravascular medical products without using animal or human clinical studies for this data. Such intravascular situations include determining properties such as drag and steerability of catheters, recoil of vascular stents, and clinician training. In fatigue testing, the latex mock arteries are used to receive deployed products and are then repeatedly pressurized at biologically relevant pressures to determine the long term durability of the product. By matching dimensions and pressure-volume relationships (compliance) of these latex tubes, researchers have a reliable means to evaluate and predict product lifetimes. The problem with latex mock arteries is two-fold: First, they are opaque so the product inside the artery cannot be seen during evaluation of the integrity of the product or during clinical training sessions. Second, latex tubes fatigue; therefore, the loading that they place on the internalized products varies with time. During long term durability studies, latex tubes may have to be replaced as often as every 100 million cycles. This can be problematic with products that are difficult to redeploy. We have developed a clear silicone mock artery system that allows us to fabricate three-dimensional objects, including tubes with precise geometric and mechanical properties. Our evaluations show that the mock arteries can be stressed up to 400 million cycles with little or no change in mechanical properties. We are in the process of continuing evaluations to determine long term durability.

Keywords:

mock, arteries, stents, in-vitro, durability, accelerated, fatigue, compliance, silicone, latex.

Introduction

During the last century or so the ability of the human race to heal its own medical problems has experienced a growth rate that staggers the imagination. Not only have the technologies and techniques that are available to the physician undergone remarkable change, but also the ability of the scientists and engineers to develop testing protocols to determine the safety and efficacy of devices, diagnostic techniques, and pharmaceuticals has increased dramatically.

Just a hundred years ago most new medical approaches were experimentally evaluated on patients that were either unaware of the consequences of the experimentation, or were unaware that the experimentation was even occurring. As approaches to medical experimentation became more civilized and controlled, the gradual introduction of experimental animals into the medical developer’s repertoire occurred. Although ethically better than using humans in early experimentation, animals offered many complications, including hematologic and immunological systems that were not directly comparable to humans, issues concerning costs and, more recently, animal rights. Another important consideration in the development of medical products, particularly targeted for the cardiovascular system, is the need to generate valid data from experiments at an accelerated rate so that ten years of implantation durability can be determined in a shorter period of time. This has brought us to a point wherever more experimentation is done with instrumentation designed to replicate the mechanical, chemical, or physical environments that the medical products will experience when implanted. These type of evaluations are usually referred to as in vitro experiments. When dealing with devices that are targeted for the continuously pulsing cardiovascular system, such as artificial hearts, left ventricular assist devices, stents, vascular grafts, heart valves, septal defect closure devices, intra-aortic balloons, etc., the issue of dynamic compliance in these in vitro experiments becomes extremely important.

Radial compliance of a tube is defined as the relationship between the radius and the internal pressure of that tube. Although there is general agreement on the definition of compliance, there have been a variety of techniques used to evaluate this parameter. Since most of the initial interest in the radial compliance of tubes was generated because of a concern relating to the long-term patency of vascular grafts [1-6], it has been postulated that it is the internal change in radius that most significantly affects the blood flow disturbance or the interaction of a graft or synthetic tube with an indwelling stent or catheter. A review of the literature indicates large variations, not only in the estimation of in vivo compliance, but also in the experimental designs used to evaluate compliance of both natural vessels and vascular grafts. Some of the work was done on excised vessels [7-8] or natural vessels that were exposed to air [9-10]. Many of the previous techniques utilized X-rays, and some took advantage of cantilevered beams [11-13] touching the vessel or vascular graft in order to determine the change in external diameter versus pressure. Additional experiments were reported using various video techniques [14-19]. More recent investigations evaluating mechanical compliance found that the three-dimensional compliance of vascular grafts and mock latex arteries was dependent upon pulsation rate or testing frequency [20-22]. This dynamic radial compliance directly affects the design of reliable protocols. There is, of course, a desire to acquire information about the durability or fatigue resistance of compliant products as quickly as possible. Consideration of the frequency response of the product under test as well as the time dependent characteristics of other experimental materials used, such as synthetic arteries housing vascular stents, becomes extremely important. Accelerated testing is plagued by problems, not just in the mechanical engineering area, but also in metal fatigue analysis and radiation of polymers. For example, increasing the dose of radiation or the storage temperature are common types of accelerated testing. While informative, these types of experiments, as well as all accelerated tests, are never perfect in design or concept. The reason, of course, is that high levels of exposure to an experimental variable, whether it is temperature, radiation, or pressure versus time, can never exactly replicate the real time situation. Experts in accelerated testing know this and take special precautions to evaluate what impact the unnatural stressing or loading has on the product’s ability to respond.

In mechanical testing particularly, it is important to spend sufficient time on initial evaluation of the frequency response of all components before finalizing test protocols. As has been shown in vascular graft and mock artery testing at high frequencies, one needs to be concerned with how the motion of the graft or the mock artery containing the stent or catheter responds to increasing rates of testing [21-22]. The key to understanding these frequency dependent changes lies with consideration of the materials as well as the limitations of the instrumentation used during the accelerated testing. It is clearly the goal of long-term durability testing to evaluate properly the fatigue resistance of the product, and to do it in a time frame that is as short as possible. This assists early detection of design flaws, the schedule of product development personnel, and the rapid submission of products to regulatory agencies.

The general goal of accelerated testing should be to neither overload nor underload the product. Overloading can lead to unpredicted and unjustified product failure and may result in the tendency to ignore failures that occur, with potentially disastrous results. On the other hand, understressed products induce unjustified confidence in the long-term durability of the devices and may result in in vivo failures with equally disastrous results. A recent paper indicated some of the errors to avoid when monitoring the high speed testing of vascular grafts or synthetic arteries that contain medical products such as stents, stent grafts or catheters [23, 24].

Although several of the aforementioned references have dealt with issues concerning the frequency dependent properties of mock vessels, there has been a more recent reference to problems associated with using latex in long term in-vitro studies [25]. This paper examined the tendency for high precision latex mock arteries to change their mechanical properties when tested at various high frequencies and pressure differentials in a 37° C distilled water environment. Table 1, taken from that reference, shows the dynamic internal radial compliance of several tubes whose initial internal compliances were measured at 72 bpm over a pressure differential of 180/80mm Hg, and then were tested at 1000 bpm at the same pressures. These tubes were then mounted on an accelerated durability tester and tested at 1000 bpm at 180/80mm Hg. The dynamic internal radial compliance was re-evaluated at approximately 102 and 180 million cycles of pulsation. After 180 million cycles, the low frequency dynamic radial compliance of these vessels had nearly tripled, while the dynamic radial compliance at high frequency showed an increase of around 100%.

Table 2 shows a second experiment, with dynamic internal compliances measured at 72 and 1600bpm. Note a 100% to 200% increase in compliance during these 180 million cycles of testing.

This magnitude of change in compliance is unacceptable for long term in vitro experiments. Using new tubes every 50-100 million cycles would be the only acceptable approach.

The purpose of the work described below was to evaluate a new material that would allow testing of experimental medical products deployed in mock bodily parts, including ventricles, arteries, etc., and tested for a least a period of 400 million cycles, which would represent an in vivo implantation period in humans of about 10 years. An additional consideration was to have an optically clear material for visual inspection without explantation or internal visualization devices such as boroscopes.

Materials and Methods

Since it was our intention to develop a new material for the fabrication of mock arteries and organs that had superior long term durability and clarity to natural latex rubber, our first task was to obtain a large number of silicone and polyurethane candidates. We initially screened these for clarity when cast at thicknesses close to what we predicted would be necessary for the production of mock vessels and organs with biologically relevant internal compliances. It became obvious that most of the materials had optical and or hardness properties that caused their elimination from the list of potential candidates. In the long run we found that certain noble metal catalysts incorporated into specific prepolymers resulted in materials that were virtually optically clear and yielded wall thickness versus compliance properties that were usable in an experimental scenario. These silicone materials are used via a dipping process by reducing the solids content and viscosity with xylene. These materials also require several rotational and pre-curing steps before they can be put into the final curing oven. These steps are necessary to insure the optical clarity of devices up to wall thicknesses of nearly 5.0mm. The first experiment involved the dipping of precision ground rods into the uncured silicone material to produce twelve tubes with various wall thicknesses. These twelve tubes were then, one at a time, mounted on a computer-controlled dynamic internal compliance tester. Their compliances were evaluated at 72bpm, with a pressure differential of 180/80mm Hg, and also evaluated at 1000 bpm with the same pressure differential of 180/80mm Hg. These twelve tubes were then mounted on a dynamic durability testing device that allows for the repeated pulsing and expansion of tubing. The experimental conditions included internal and external solutions of distilled water that were kept at 37° C throughout the testing. The tubes were pulsed at 1000 bpm at a pressure range of 180/80mm Hg per cycle.

Periodically throughout the testing, some or all of the tubes were removed to determine if the dynamic internal compliance of the tubes was changing as a result of the pulsatile stresses, similar to the experiment referred to in reference 25. After these intermediate compliance tests were done, the tubes were remounted on the accelerated tester and the pulsations continued. This testing continued through 400 million cycles of pulsatile stressing. At the end of the experiment all tubes were removed and again their dynamic internal compliances evaluated.

A second experiment was initiated in which the tubes were put through a double final curing step, so as to evaluate the influence of this extreme curing process on any changes and long term compliance of the tubing. In this case, however, nine tubes of approximately the same compliance were produced and were subjected to the exact same experimental conditions as previously described. A difference in the protocol was the determination of the internal diameter and wall thickness.

Results

Table 3 is a compilation of the results obtained from the initial twelve experimental tubes. The dynamic internal compliance was evaluated at 72 bpm at a pressure range of 180/80mm Hg. Note no increase in compliance as seen previously with latex tubing. In fact, there is a reduction in compliance, indicating a strengthening of the tubes with testing up to 400 million cycles.


Table 4 shows the results from the evaluation of the same twelve tubes, only the compliance in this case was determined at a frequency of 1000 bpm with the same pressure range. There was no indication that the tubes were weakening, as would be evidenced by an increase in compliance.

The second part of the experiment evaluated tubes that had additional levels of final curing. Table 5 shows the results so far when evaluating the compliance at 72 bpm or 1100 bpm. The initial compliance of all nine tubes were close to each other and the final compliance after 100 million cycles of testing did show some reduction in compliance, but substantially less that had been seen at the lower cure time.

Table 6 shows the evaluations of inside diameter and wall thicknesses of the same tubes, which show a small and probably inconsequential change in diameter in these tubes.

Conclusion

The utilization of noble metal catalyzed silicone polymers in a dipping process to produce mock arteries has resulted in tubes that show an elimination of the aging related changes in compliance that has been seen with the standard latex tubing used for mock arteries. This dramatic change in long term physical properties, coupled with the near optical clarity of the tubes, indicates the superiority of this material for the production of mock arteries and other organs. They are ideal for long term durability evaluation of medical products, short term training, or evaluation that requires superior optical clarity.

References

1. Kinley, C. E., and Marble, A. E., Compliance. J. Cardiovasc. Surg., 21,
163-170 (1980).

2. Clark , R. E., Apostolou, S., and Kardos, J. L., Mismatch of Mechanical
Properties as a Cause of Arterial Prosthesis Thrombosis. Surg. Forum, 27,
208-217 (1976).

3. Gonza, E. R., Mason, W. F., Marble, A. E., and others, Necessity for
Elastic Properties in Synthetic Arterial Grafts. Can. J. Surg., 17, 176-179 (1974).

4. Lyman, D. J., Fazzio, F. J., Voorhees, H., and others, Compliance as a
Factor Effecting the Patency of Copolyurethane Vascular Grafts. J.
Biomed. Master. Res., 12, 337-339 (1978).

5. Abbott, W. M., and Bouchier-Hayes, D. J., The Role of Mechanical Properties
in Graft Design. In: Graft Materials in Vascular Surgery. Edited by H. Dardik.
Miami: Symposia Specialists, Inc., 59 (1978).

6. Baird, R. N., Kidson, I. G., L’Italien, J. G., and Abbott, M. W., Dynamic
Compliance of Arterial Grafts. Am. J. Physiol., 2, 568-572 (1977).

7. Bergel, D. H., The Static Elastic Properties of the Arterial Wall. J.
Physiol. (London), 156, 445 (1961).

8. Bergel, D. H., The Dynamic Elastic Properties of the Arterial Wall. J.
Physiol. (London), 156, 448 (1961).

9. Patel, D. J., De Freitan, F. M., Freenfield, J. R., and Fry, D. L., Relation of
Radius to Pressure Along the Aorta in Living Dogs. J. Appl. Physiol, 18,
1111 (1963).

10. Paterson, L. H. Jensen, R. E. and Parnell, J., Mechanical Properties of Arteries
In Vivo. Circulation Res., 8, 622 (1960).

11. Walden, R., L’ Italien, G. J., Megerman, J., and Abbott, W. M., Matched
Elastic Properties and Successful Arterial Grafting. Arch. Surg., 115,
1166-1169 (1980).

12. Gow, B. S., and Taylor, G., Measurement of Viscoelastic Properties of Arteries
in the Living Dog, Circulatory Research, 23, 111 (1968).

13. Gow, B. S., An Electrical Caliper for the Measurement of Pulsatile
Arterial Diameter In Vivo. J. Appl. Physiol, 21, 1122 (1960).

14. Gotoh, F., Muramatsu, F., Fukuuchi, Y., Okaysu, H., Tanaka, K., Suzuki, N.,
and Kobari, M., Video Camera Method for Simultaneous Measurement of
Blood Flow Velocity and Pial Vessel Diameter. J. Cereb. Blood Flow Metab.,
2, 421-428 (1982).

15. Assmann, R. and Henrich, H., A Video-Angiometer for Simultaneous
and Continuous Measurement of Inner and Outer Vessel Diameters.
Pflugers Arch., 376, 263-266 (1978).

16. Halpern, W., Mongeon, S. A., and Root, D. T., Stress, Tension and
Myogenic Aspects of Small Isolated Extraparenchymal Rat Arteries. In:
Smooth Muscle Contraction, edited by N.L. Stephens. New York : Dekker,
427-456 (1984).

17. Intaglietta, M. and Tompkins, W. R., On-Line Measurement of
Microvascular Dimensions by Television Microscopy. J. Appl. Physiol., 32,
546-551 (1972).

18. Wiederhielm, C. A., Continuous Recording of Arteriolar Dimensions with
a Television Microscope. J. Appl. Physiol., 18, 1041-1042 (1963).

19. Halpern, W., Osol, G., Coy, G. S., Mechanical Behavior of Pressurized In
Vitro Prearteriolar Vessels Determined with a Video System. Annals of
Biomedical Engineering, 12, 463-479 (1984).

20. Conti, J. C., Strope, E. R., Rohde, D. R. and Greisler, H. P., A New Technique
to Determine Vascular Compliance In Vivo, National Heart Lung and
Blood Institute Contractors Meeting. Louisville , KY , 1989.

21. Bozzi, R., Conti, J. C., Soldani, G., Spence, L., Strope, E., and Withrow,
D., Relating the Frequency Dependent Radial Compliance to the Tensile
Modulus of Polyurethane and Latex Vascular Grafts. Transactions 5th
World Biomaterials Congress, 434, U2 (1996).

22. Conti, J. C., Strope, E. R., Rohde, D. J., and Spence, L. D., Frequency
Dependent Radial Compliance of Latex Tubing. Biomedical
Sciences Instrumentation, 33, 524-529 (1997).

23. Conti, J. C., Strope, E. R., and Price, K. S., Sources of Error in Monitoring
High Speed Testing of Vascular Grafts, Biomedical Sciences Instrumentation,
34, 240-245 (1998).

24. Conti, J. C., Strope, E. R., Price, K. S., and Goldenberg, L. M, The
High Frequency Testing of Vascular Grafts and Vascular Stents: Influence
of Sample Dimensions on Maximum Allowable Frequency. Biomedical
Sciences Instrumentation, 35, 339-346 (1999).

25. Conti, J. C., Strope, E. R., and Price, K. S., High Frequency Testing of Grafts
and Mock Arteries for Stents: Influence of Testing Frequency on
Durability, Transactions Society of Biomaterials, 403 (1999).