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Infarct expansion and extension of the border zone play a key role in the progression of heart failure after myocardial infarction. Increased wall stress, along with complex cellular and extracellular changes in the surviving myocardium, underlie these events and contributes to the adverse cardiac remodeling that drives ventricular dilation and progression of heart failure. Recently, there has been much interest in the development of biopolymers that can be injected into the infarcted myocardium in order to increase its stiffness and thus reduce mechanical stress on the surrounding myocardium. Here we discuss the findings of recent animal studies that have noted improvements in contractile function or cardiac remodeling using either natural or synthetic biomaterials, as well as several that did not. Besides offering physical support to the injured myocardium, injectable biomaterials could also serve the purpose of fostering cardiac repair by functioning as a protective scaffold for stem cell or drug delivery.
Almost 3 decades ago, myocardial infarction (MI) was associated with very high morbidity and mortality. If a patient survived an acute infarct their prognosis was very poor and highly dependent on post-injury complications (1). Acute complications such as rupture of the myocardium, contractile dysfunction, valvular disease, and heart block or long-term complications such as heart failure, and arrhythmias severely limited survival. Marked improvements were not seen until the mid-80s with the development of interventional and pharmacological therapies and the focus on early coronary reperfusion (1). With clot-busting agents, coronary artery bypass grafting (CABG), percutaneous coronary intervention (PCI), and long-term treatment with ACE inhibitors and β-blockers, survival has increased to > 80% for the 30 day period post-MI; although, outcome is highly dependent on patient risk level and hospital performance (2,3).
Of course, not all patients exert the same level of response or have the same infarct size. In addition, high risk patients are more prone to infarct expansion and border zone extension due, in part, to increased mechanical wall stress placed on the injured myocardium, which inevitably progresses into heart failure (4,5). The border zone is characterized by a hypocontractile myocardium that is nonetheless perfused. With time adverse cellular and tissue remodeling events contribute, along with mechanical forces, to infarct expansion and border zone extension. This progressive maladaptive response is associated with increased myocyte apoptosis/necrosis and slippage, extracellular matrix (ECM) disturbances marked by fibrosis and changes in collagen makeup, and a complex wound healing process (6,7). In time, the heart loses most of its function and pumping capacity due to chamber dilation and thinning of the ventricular wall (8). Despite extensive research and discovery of large number of biomarkers that represent possible therapeutic targets, how the heart reaches this irreversible stage is still unclear.
Intense research has been focused recently on late stages of heart failure and novel strategies such as regenerative stem cell therapy and tissue engineering have emerged (9–11). Some modest improvements in cardiac function have been documented in clinical trials of stem cell treatment (12); however, reversing adverse remodeling and regenerating a fully functional and integrated myocardium is still a dream. Alternative strategies aimed at preventing adverse remodeling may represent a much more attainable goal. With that in mind, some groups have focused efforts on devising ways of preventing the initial injury, especially in patients with high risk factors, while others have focused on preventing adverse remodeling after the injury has occurred (Fig. 1).
In this review we focus on therapeutic strategies that prevent adverse remodeling through the use of acellular biomaterials injected directly into the infarcted area of the heart in the early stages post-MI. The basis for this strategy is to decrease compliance of the infarcted myocardium, i.e., increase infarct stiffness, and thereby lessen wall stress on the surrounding myocardium. A dramatic proof of concept supporting the utility of this approach was recently reported (4). A dermal filler agent sold to treat facial wrinkles was injected into a sheep anteroapical infarct 3 h post-MI (Fig. 2). In addition to increased infarct thickness and stiffness, they reported attenuated left ventricular remodeling and improved global LV function assessed 8 weeks later (Fig. 3).
MI prevention therapy has saved hundreds of thousands of lives but today a large number of patients are still seen in the ER with acute MI, which is responsible for more than three quarter of heart failure cases in the USA alone (5). The smaller the infarct and faster the interventional procedure to reperfuse the myocardium, the higher the rate of survival and long-term prognosis, and less likely the patient is to suffer from acute complications or permanent damage. After reperfusion, pharmacological interventions play an important role in the acute and long-term treatment, to improve systolic performance, decrease arrhythmias, and reduce infarct expansion by decreasing cardiac workload. In some cases, e.g., severe multivessel coronary artery disease, invasive procedures such as CABG are unavoidable. However, no treatment or procedure prevents disease progression into the more debilitating stages of heart failure, especially in patients with high risk factors such as age, diabetes, PVD and CAD (14). Today more patients survive an MI, resulting in many more patients developing heart failure.
The early progression of cardiac remodeling post-MI is easily understood from Laplace’s Law (T = PR/h), where wall stress in the left ventricle (T) is equal to the pressure in the chamber times radius of curvature (R) divided by thickness of the myocardial wall (h) (13,15). While ischemia and altered cellular function are responsible for initial heart damage, initiation of the myocardium into a progressively dysfunctional state that foretells heart failure is due to alterations in mechanical and structural properties (16–19). Loss of pumping efficiency at the onset of MI due to ischemia will elevate the amount of the blood in the ventricles and increase chamber pressure and LV dilation, eventually shifting the heart structure from a normal elongated shape to a more spherical shape. Although the initial response to stretching of normal cardiac myocytes is an increase in contractility to maintain cardiac output, this compensatory (Frank-Starling) mechanism is lost with overstretch, which will further increase the hemodynamic load and reduce LV ejection fraction (20). With MI, dilation of the heart and wall thinning increase wall stress, and wall stress increases chamber dilation and wall thinning, setting up a vicious cycle.
Many clinically effective but highly invasive approaches have been devised to restrain cardiac dilation and thinning of the ventricular wall (21,22). Approaches include surgeries, such as the now discredited Batista procedure and the Dor procedure that reshape the dilated spherical heart into a more elongated natural form to restore pumping capacity (23,24), or implanting devices such as wraps that provide physical support and reduce wall stress (25–30). These effective but highly invasive techniques have not been adopted in early stages post-MI.
There is growing interest in developing techniques that are effective but minimally invasive and can be applied in the early stages post-MI. One approach is the delivery of natural and/or synthetic materials (hydrogels) percutaneously into the injured myocardium by catheter via an epicardial approach or by intracoronary injection (31,32). Studies in small animals such as mice, rats and rabbits (33–42) and larger animals such as sheep and swine have been performed (32,43). Various types of hydrogels with different volumes and concentrations have been injected at different times post-MI, ranging from onset of infarction to 8 weeks, and followed over a period of time ranging from 2 to 16 weeks (31–43). Regardless of the differences in models and methodologies, there is a consensus that there is improvement in the myocardium by reducing hemodynamic stress, thickening of the ventricular wall, and limiting maladaptive cardiac remodeling of MI.
Biomaterials such as polymers were originally seen as a scaffold for delivery of cells or molecules into injured tissue. Researchers who pioneered these procedures noticed that control animals who received injection of a bio-inert material alone exhibited limited progression of ventricular remodeling into more deleterious cardiac dysfunction. One benefit of the filling agent was shown to be mechanical support thereby preserving wall thickness and limiting infarct expansion and border zone extension (31–43). In addition to providing physical support, polymers, especially the natural ones, exhibit bioactivity, which may stimulate angiogenesis and progenitor cell infiltration with degradation (44,45). A third possible beneficial action of acellular biomaterials may be the trapping of necrotic and apoptotic cells of the infarcted area into a rigid polymerized scaffold that limits release of pro-inflammatory danger-associated molecular patterns (DAMPS) into surrounding tissue (Fig. 4). This mechanism would decrease surrounding molecular stress, which is highly correlated with cardiomyocyte death and contractile dysfunction (44), and leads to infarct spread, border zone extension, and heart failure. Different parameters such as polymer crosslink density, molecular affinity, degradation time, and in situ liquid to solid transition time play a key role in modulating the outcomes of the hydrogel in use, such as support, bioactive material binding, cellular infiltration and angiogenesis (45).
Many different types of biomaterials have been tested in different animal models: natural polymers or hydrogels, which undergo controlled liquid-to-solid transition, composed of fibrin (fibrinogen and thrombin), collagen, alginate gel, matrigel, or hyaluronic acid and chitosan, and synthetic ones such as αCD-MPEG-PCL-MPEG, Dex-PCL-HEMA/PNIPAAm and poly (NIPAAm-co-AAc-co-HEMAPTMC) (31–43). Even though these compounds act as bulking agents they differ in their chemistry, extent of polymerization, porosity, hydrophilicity, stability, elasticity, and biodegradation (45). A summary of injectable hydrogels and their assessment in animal models of MI can be found elsewhere (15).
One major advantage of natural polymers is their ability to degrade with time, making them a good source of bioactive material in addition to their structural enhancement action. Chrisman et al. and Huang et al. assessed the consequences of injecting fibrin glue 1-week post-LAD ligation/reperfusion in a rat model (34,37,42). Both observed mechanical and structural improvements after 5 weeks as evidenced by preserved infarct thickness and LV fractional shortening, as well as neovascular formation likely related to the bioactive effect of fibrin. Yu et al. reported comparable results using a similar model studied over 10 weeks post-MI, but in this case fibrin glue was delivered 5 weeks post-MI, thus allowing for infarct remodeling to be highly established at the time of treatment (41).
Dai et al. tested the effect of injecting collagen-based hydrogel into the infarct area of a rat heart (33). While mechanical and functional improvements were noted, no significant angiogenesis or cell infiltration was found. In contrast, Huang et al. reported that collagen increased capillary density and myofibroblast infiltration into the infarct area of their ischemia-reperfused rat model (37). These differences may be due to differences in the model used (permanent ischemia versus ischemia/reperfusion) as well as the concentration, volume, and type of collagen administered. Other collagen-based thermal hydrogels known as matrigel have been tested (35,37,46). Matrigels also contain basement membrane components that boost bioactivity. In both rat and mouse MI models, matrigels improved cardiac function and increased scar thickness (35,46). Capillary density was also significantly higher in the rat reperfusion MI (37).
An interesting widely-studied bulking polymer is in situ-forming calcium-cross linked alginate. By introducing a tripeptide known as RGD (arginine-glycine-aspartate) or a pentapeptide YIGSR (tyrosine-isoleucine-glycine-serine-arginine) alginate can be modified to promote cell binding and stability after injection (38,47). Modified and non-modified alginate have been studied in both swine and rat models (31,32,36,38,47). In a non-reperfused MI rat model, Tsur-Gang et al. studied progression of cardiac remodeling over 9 weeks in response to alginate injection 7 days post-MI (38). While both alginate types showed significant and similar increases in scar thickness and angiogenesis, the effect of modified alginate on the progression of LV dilatation and dysfunction was reduced and similar to the controls compared to the non-modified alginate. Positive effects of non-modified alginate on cardiac function have been also demonstrated by Landa et al. who tested the same infarct model over the same period of time (36). In a LAD ligation/reperfusion rat model, Yu et al. studied the effect of both alginate types after their delivery 5 weeks post-MI over a period of 5 weeks (47). In this case, there was significant cardiac functional improvement and wall thickness over the controls with both types of alginate; moreover, higher arteriole density was observed with modified alginate. This might be explained by the bioactivity of adhesive alginate and its capacity to interact with up-regulated sites in the reperfused myocardium promoting angiogenesis. Reperfused myocardium also has a higher capacity for angiogenesis (48).
Mukherjee et al. performed similar experiments in a swine model but with a combination of both fibrin and non-modified alginate delivered 1 week post-obtuse marginal 1(OM1) and OM2 ligation for a period of 3 weeks (31). The increase in wall thickness and regional MI stiffness, and reduction in scar size was accompanied by an improvement in cardiac index compared to infarcted hearts receiving saline. Leor et al. who delivered non-modified alginate into the infarcted myocardium via an intracoronary catheter injection directly into the infarct-related coronary artery of swine 4 days after a transient balloon occlusion and monitored for 8 weeks, reported similar data (32). MI-induced LV enlargement was either prevented or reversed. While in most studies the MI is induced after thoracotomy and the biomaterial is delivered directly into the epicardium, the approach of Leor et al. was minimally invasive and more clinically relevant especially as the animals were pre-medicated with aspirin, clopidogrel, lopressor, and amiodarone and aspirin was continued throughout the study.
Other researchers studied the effect of hyaluronic acid (HA) on progression of myocardium remodeling. HA is an abundant polysaccharide in the body with positive effects on wound healing (49–52). After mixing with a thiol-terminated PEG cross linker, this polysaccharide was delivered to a rat MI model two weeks after infarction. After 4 weeks, LV remodeling was limited with a significant decrease in infarct size, which was associated with an increase in cardiac function and capillary density even after complete polymer degradation (40). Similar results were seen in an ovine infarct model using methacrylated HA (MeHA) (43). Differences in the mechanical properties of the hydrogel likely contribute to outcome (43). Ifkovits et al. demonstrated that high modulus MeHA had higher and better outcome on LV progression and cardiac function versus low modulus type, highlighting the impact of hydrogel modality and mechanical properties on its remodeling actions (43).
Chitosan is a polysaccharide-based thermal hydrogel widely used in many tissue engineering applications (53–55). Chitosan injected one week after infarction in a MI rat model induced a significant increase in cardiac function and scar thickness after 4 weeks, as well as angiogenesis even after significant degradation (56). Thus, this is another report suggesting the importance of the polymer bioactivity in addition to its supportive mechanical function in repairing the heart.
Several types of synthetic hydrogels have been developed and tested. Their relative lack of bioactivity allows the study of the contribution of structural improvement by itself to ventricular remodeling progression after an infarction. αCD-MPEG-PCL-MPEG is a poly (ethylene glycol) plus α-cyclodextrin self-assembly degradable polymer that has been tested in a non-reperfused LAD ligated model in both rats and rabbits (57,58). In the rat model, the hydrogel was injected 5 min after infarction and the impact on the heart assessed after 4 weeks (57). Despite decreased infarct size and improved cardiac function, neovascularization was not observed. Similar results were obtained by Jiang et al. who injected the polymer 1 week after LAD ligation in a rabbit model for the same period of time (58).
Another dextran-based degradable thermal polymer is Dex-PCL-HEMA/PNIPAAm (40). This material was delivered to a rabbit heart 4 days after ligation of the proximal left coronary artery (40). Functional improvement and wall thickness were documented 4 weeks later even with complete polymer degradation. A similar hydrogel named poly (NIPAAm-co-AAc-co-HEMAPTMC) was synthesized by Fujimoto et al. and tested in the ligated LAD rat model (59). The material was delivered two weeks after infarction and significantly improved cardiac function and LV wall thickness were seen 8 weeks later compared to the controls.
Some researchers have focused on the long-term impact of non-degradable and non-bioactive hydrogels on LV adverse remodeling progression post MI. If the structural support by itself is enough to decrease wall stress by decreasing hemodynamic load than the biomaterial should improve and maintain the cardiac function over a long period of time. However, both Rane et al. and Dobner et al. report results that refute this idea (60,61). In the study of Dobner et al., a non-degradable vinyl sulfone poly(ethylene glycol), PEG-VS, was used. The hydrogel was injected 2 minutes after LAD ligation in a rat model. Analyses were performed at both 4 and 13 weeks post-injection. Despite the increase in wall thickness throughout the study, significant cardiac improvement was only maintained the first 4 weeks and was lost after 13 weeks (60). Similarly, Rane et al. reported identical results using non-degradable PEG-ASG hydrogel. This material was delivered 1 week post-MI in a non-reperfused rat model for a period of 6 weeks. Even though wall thickness was significantly higher in treated animals, magnetic resonance images (MRI) showed non-significant differences in the negative LV remodeling with a decrease in cardiac function and expansion (61). These findings would suggest that acellular injections might find utility as a bridge therapy allowing time for cardiac regenerative approaches to take hold.
While overall the results from animal studies on the effectiveness of acelular biomaterials to limit infarct expansion seem promising, the negative studies highlight the need for a more systematic approach to define the optimal material properties of the injected polymers. Of course, the animal models used in these studies do not resemble the clinical model in many respects. However, it is worth noting that the clinical situation involves long-term pharmacological therapy designed to decrease stress on the heart. This by itself could be an important factor in making biomaterial injection a good supportive therapy with perhaps a better prognosis than the one in animals where the MI is induced with no subsequent drug treatment. Along these lines, hydrogel delivery could be combined with stem cell regenerative therapy and/or used as a depot for drug delivery (62). Decreasing wall stress and cellular stress, while increasing angiogenesis, may also be a better environment to nurture stem or progenitor cells to promote regeneration of lost myocardium.
This work was supported by a grant from NHLBI to GWB (5R01HL088101-05).