It is vital that any product performs to the standards set by the manufacturer’s claims. Many medical devices are born from the desire to satisfy a previously unmet clinical need. To be successful, a new product must elicit this functionality, at or above the currently available solutions and techniques. The functionality of a cardiac device also encompasses its ability to perform the desired task without compromising any other biological process. For example, a transcatheter-delivered aortic valve replacement should not compromise the function of the mitral valve or the vasculature through which it is delivered to the heart. These specifications are often outlined by regulatory committees such as the International Organization for Standardization (ISO), as seen in Table 16.2 describing the operational specifications for cardiac valve prostheses.

For implantable cardiac devices such as transcatheter and surgically implanted valves, the question of functionality typically includes the pathway in which the device is delivered. For example, surgically implanted devices require a suitable method of fixation to the heart, such as the sewing ring of a prosthetic heart valve. However, these parameters can be greatly complicated when one chooses to deliver such devices through minimally invasive techniques, such as through the wall of the heart (typically the apex) or through the vasculature without placing the patient on cardiopulmonary bypass. Hence, not only must the delivery system be able to position the device in its intended anatomical location, but additionally it must not induce any adverse effects on the patient. For instance, the patient populations in need of aortic valve replacements will often present with calcified aortic arches. A calcified aorta can be problematic when delivering a device from the femoral artery, since the delivery system must not dislodge the calcifications nor puncture the aorta, both potentially fatal incidents.

Another excellent example of device functionality involves the development of transcatheter-delivered heart valves to treat patients with highly stenotic aortic valves. Note that these technologies are described at length in Chap. 10, along with a discussion of the potential benefits and detriments of device functionality with regard to both the implanting physician and the patient. To date, companies developing these new transcatheter-delivered technologies have been very specific in their description of the patient populations for which the valves are designed, as well as the conditions under which the valves should be implanted. As such, they create extensive documentation of instructions for use (IFU). By doing so, they develop guidelines for acceptable usage of the device and attempt to ensure that the devices are used only in situations that correspond to the conditions under which the device was critically designed to function. Although device manufacturers set guidelines in the IFU regarding how the device is to be utilized, in rare instances a physician may choose to use it otherwise (e.g. use an aortic valve device to replace the pulmonary valve). This action is called off–label use. Off-label use is something that does occur and, in many instances, has saved numerous lives. However, even if a physician has used the device for a particular purpose outside of the IFU, in no way can the device manufacturer encourage or advertise this use unless it has gone through animal and clinical testing for that particular use (see Chap. 17 for additional details regarding preclincial testing).

Biocompatibility related to medical devices is generally defined as the ability of a material to perform with an appropriate host response in a specific application [7].

In 1987 Williams described how the biocompatibility of a material can be qualitatively evaluated to assess its relative performance when implanted. The succinctness of this term is often misleading when applying the principles of biocompatibility to cardiac devices, due to the large number of different materials that are commonly used to manufacture such modern devices. For example, tilting disk valve prostheses such as the St. Jude Masters series (St. Jude Medical, MN; Fig. 16.2) are manufactured from a variety of materials including pyrolytic carbon for the disks and Dacron® for the sewing cuff. One needs to consider that each of these materials may elicit different responses from the host body. Hence, they need to be individually assessed and quantified before the cardiac device is implanted. Examples of such tissue responses can be seen in Table 16.3.

The mechanisms related to how the host or patient may respond to different components of a cardiac device must be extensively evaluated to ensure that the correct materials are selected for the final device design. This evaluation can be complicated by the fact that it is hard to replicate, in vitro, the range of human immune responses that may exist. To address this, all cardiac devices must also undergo rigorous animal testing, yet it is important to note that this too can sometimes provide misleading results on species-specific bioreactivities. An example of this can be seen in the design and development of the Braunwald-Cutter heart valve ball and cage prostheses (Fig. 16.3), whereby cloth-covered cage struts were designed to encourage endothelialization and subsequently decrease any chance of thrombolytic events [9]. Extensive testing in the mitral position of pigs, sheep, and calves showed promising results, thus the device was approved for human clinical trials. However, the device did not elicit the same host responses when implanted in humans; rather it resulted in aggravated wear on the cloth cage struts and, more critically, the formation of debris embolization. It should also be emphasized that host responses are not limited to immune reactions. For example, implanted tissue heart valves remain susceptible to accelerated prostheses leaflet calcification. More specifically, it has been reported that this type of calcification is initiated by reactions between the extracellular fluid and the leaflet membranes, creating calcium phosphate mineral deposits [10]. As a result, much research is ongoing to enhance leaflet materials and minimize bioreactivities, including calcification inhibitors on the valves that limit mineral deposition on the implanted materials. Furthermore, a device’s material can also display unexpected interactions with the host, as evidenced by the earliest versions of the Starr-Edwards caged-ball mechanical heart valve. The valve design used a silicone ball that was found to absorb lipids from the blood and thus swell [11]. In addition to resultant poor valve function, this also caused the silicone balls to become brittle; in turn, this increased the possibility of ball fracture and consequent embolization of small fragments into the arteries downstream of the valve position.

As well as exhibiting the appropriate level of biocompatibility, any device implanted into a human heart must also display sufficient durability and longevity. It should be emphasized that this poses a unique problem in the field of implantable cardiac devices, due to the extremely large number of cycles that such prostheses are expected to endure. Furthermore, there are potentially life-threatening implications due to device failures. One needs to consider that the heart is itself a highly dynamic organ which can elicit three to fourfold increases in cardiac output (internal pressures); hence, the mechanical stresses on an implanted device can increase four to sixfold during exercise.

The relative durability of any cardiac device is further challenged by directly interacting with the turbulent and pulsatile flows of blood, which is the case for prosthetic cardiac valves. More specifically, the flow of blood through mechanical prostheses, particularly tilting disk valves, can lead to a particular type of erosion called cavitation whereby gaseous bubbles form and violently collapse on the surface of the valve disks due to the sudden temporary decrease in pressure during closure. This phenomenon has resulted in fractures of the valve disks and was originally observed in the 1980s, in association with the failure of the Edwards-Duromedics valve (Edwards, CA) [12]. Subsequent tilting disk valves, such as the St. Jude bileaflet valve, have been designed to critically reduce this particular material erosion by reducing the area of the tilting disks.