1. Introductıon
The field of engineering has been accustomed to dealing with designing materials at their limits for properties including yield stress, endurance limit, and rupture life according to the use of material and the criticality of the application. After the field of biomaterials has been introduced to the field, biocompatibility of materials become one of the main design limitation for the engineers dealing with medical and biological problems.
Incompatibility is now accepted for the ultimate limit to the engineering solution of many biomedical problems. And the field of materials has been evolved into biomaterials in order to incorporate a thorough grounding in the aspect biocompatibility, along with some other considerations in the field of living systems.
In order to set a frame for the terms in the field biomaterials, European Society for Biomaterials organized a working consensus conference in 1986 and 1991. Some of the terms related to this text that gained consensus on their definitions are given below:
- Biomaterial: A material intended to interface with biological systems to evaluate, treat, augment, or replace any tissue, organ, or function of the body
- Bioactive material: A biomaterial that is designed to elicit or modulate biological activity
- Biocompatibility: The ability of a material to perform with an appropriate host response in a specific situation
- Biodegradation: the breakdown of a material mediated by a biological system
- Inherent thrombogenicity: Thrombus formation controlled by the material surface
The goal of biocompatibility is to obtain the desired effect of the device in the application for which it is intended. This includes not only optimization of desirable properties but minimization of the unwanted ones. This process starts with the selection of the material design and manufacture of the device and the testing to verify biocompatibility.
The prediction of the reaction of the device depends on the extent of the knowledge of relevant past history of the material used. This will depend on whether the actual material was used in a similar environment, a similar length of time and for a similar function. It may be necessary to extrapolate from existing data or it may be necessary to use data from chemical analogies let the knowledge of chemistry to serve as a predictor.
To have a better understanding of the context of the biocompatibility field, we need to know what the attitude of the body is to a non-self material introduced to it. Therefore, in the following chapter, we will examine the responses of host to a biomaterial.
2. Host responses to bıomaterials
2.1 The inflammatory response
Inflammation is a nonspecific physiological response to tissue damage in animals. It arises as a response to trauma, infection, intrusion of foreign materials, local cell death, or as an addition to immune or neoplastic responses. Any kind of implantation causes an inflammatory response, however the intensity of inflammation varies notably from one implant material to another.
2.2 Coagulation and hemolysis
Interaction of various blood components with biomaterials occurs immediately after implantation. It is necessary that the biomaterial does not interfere with coagulation nor with the factors that regulates it. In addition, death of blood cells, or hemolysis, by the implant should be avoided.
2.3 Adaptation
The presence of an implant, due to the implant’s chemical, physical, or electrical properties, affects the organization and elaboration of tissue elements in the vicinity, a process called adaptation or adaptive remodeling. The recognition and management of adaption are important aspects of design, development, and use of biomaterials.
2.4 Allergic foreign-body response
Body has the ability of developing a specific memory for certain non-self materials and fight with these to eliminate them. Allergy is the property of body to being especially sensitive (hypersensitive) to an agent. In some cases, body creates an allergic reaction to some implants. Therefore, the implant material should be chosen to avoid such reactions.
2.5 Chemical and foreign-body carcinogenesis
The implant material should not be a type that itself or its degradation products cause cancer. However, sometimes cancer may be developed by the host not because of the chemical properties of the implant. This happens especially when the implant has a large solid surface with little or porosity. Also the sharp ends of an implant facilitate the process.
2.6 Metal toxicity
The concentrations of all metals in the body are tightly regulated. However, the possibility that introduction of a metal from an endogenous source, that is, release of it from implant, may interfere with homeostatic regulation and produce adverse effects, either at the site of normal action of the metal, or at other sites, either directly or by interfering with other trace element-mediated processes.
2.7 Effect of degradation products on remote organ functions
Degradation products of metallic or polymeric implants may travel to and accumulate at specific remote organs like liver or kidneys and might produce adverse effects on these organs functions.
3. classes of bıomaterıals
Examining the historical development of biomaterials, Black (1999) has identified four clases of biomaterials, based on host response. These are;
- Inert: Implantable materials that elicit little or no host response.
- Interactive: Implantable materials designed to elicit specific, beneficial responses, such as ingrowth, adhesion, etc.
- Viable: Implantable materials, incorporating or attracting live cells at implantation that are treated by the host as normal tissue matrices and are actively resolved or remodeled.
- Replant: Implantable material consisting of native tissue, cultured in vitro from cells obtained previously from the specific implant patient.
According to Black, searches for inert materials are pointless. Many materials in present clinical use and those in development are interactive materials. Viable materials are the subject of active research and commercial interest. Initial examples of replant materials are in limited clinical use, while advances in control and manipulation of the genetic code in mammals suggest that no intellectual barrier exists to prevent the broad future of realization of replant materials at both the tissue and organ level. In fact, replant material (implantable, live tissue with the identical genetic code and immunological determinants of the recipient patient) represents the fulfillment of the original search for biocompatibility: implant materials demonstrating harmonious interaction.
4. methods for evaluatıng bıocompatıbılıty
4.1 Tests to evaluate inflammation
Short and long term material implantation studies are essential to understand the reaction of the host to the prosthesis. In each case, one can affirm that an implanted material will cause an inflammatory reaction in surrounding tissue. The intensity of inflammation varies notably from one material to another; nevertheless, it is recognized that destruction of the surrounding tissue accompanies inflammation.
The material in question is implanted subcutaneously to test the inflammatory activity. Depending on the duration of implantation, it is possible to study acute or chronic inflammatory response. Inflammatory response is most effectively evaluated by histological analysis through light or transmission electron microscopy. Sometimes it is desirable to obtain and analyze free inflammatory cells, which can be done by the cage technique. In this technique, the material is put in a small metallic cage and the cage is implanted subcutaneously. Then the exudates around the implant site is withdrawn and examined for inflammatory cell content. In a second technique, release of enzymes from the blood cells near the implant site is examined by immunohistochemical methods. Then the amount of enzymes released are correlated with the inflammatory response.
Chronic inflammatory response can also be analyzed by electron microscopy techniques. One can look for multinuclear giant cells with persistent neoangiogenesis and the formation of fibrous collagen and fibroblast cell capsules. Intensity of tissue reaction can be understood by analyzing the thickness of the capsule and the type of collagen present.
4.2 Infection testing
Infection of the location of implant is most commonly evaluated by histological observation of tissue-implant interface. Effectiveness of the sterilization technique is evaluated for different pathogens. A specific contaminant is introduced on the material and the amount of living microorganisms are measured after sterilization. Also, presence of leukocytes and positive result of gram staining indicates infection.
4.3 Hemocompatibility testing
In practice, a material should not cause certain events such as platelet adhesion, aggregation, blood coagulation, obstruction of blood flow for fibrin or cell deposit, and depletion of electrolytes. Therefore, it is useful to study the basic hemocompatibility of a material. A set of standard test procedures are applied to the materials and results are compared to controls.
4.3.1 In vitro hemocompatibility evaluation
In vitro methods are the first screening of the materials in question. One tests involves contact of the surface of the material with the whole blood or its components. Contact time may be predetermined or until the clotting occurs. The results should not be very short or very long for clot occurrence. In some tests, platelet-rich plasma or suspended platelets are used to determine the degree of platelet adhesion and morphology of the adherent platelets. Deposition of certain plasma proteins, including albumin, fibrinogen, or complement factors is analyzed. Protein absorption can be measured by UV-Visible, fluorescence, IR, and FTIR spectroscopy, radioimmunoassay, circular dichroism, and radiolabelling methods.
4.3.2 In vivo hemocompatibility evaluation
In vitro tests are important in predicting the hemocompatibility of materials, however, in vivo tests are still the only satisfactory methods for hemocompatibility testing. Material is implanted for a determined time period as a vascular graft, vascular catheter, arterio-venous shaft, or in a heart assist device. After removing implant, clot formation on the implant or kidney or lungs is evaluated.
4.3.3 Ex vivo hemocompatibility evaluation
A less invasive method that avoids the sacrifice of animal is ex vivo. The blood is diverted via a cannula to a chamber, or the shunt containing the test material. The blood can be returned to the animal or not. The time after which the vein is coagulated after contacting with the material is measured (Dudley clotting time).
4.4 Toxicological tests
In 1976, the US Food and Drug Administration (FDA) established a series of short and long term studies for determining the possible toxic effect of a material. Several organizations publish guidelines for testing biomaterials and devices (American Dental Association, American Society for the Testing of Materials).
4.4.1 Primary acute toxicity screening
Initial toxicity and biocompatibility screening tests are conducted in cell culture, usually of fibroblasts. Material or its extracts are put in contact with the cells. After 24 to 72 hours later, a colorant is added to medium and the dead cells are colored. Then the diameter of the colored cells from the implant material is measured and toxicity of the material is evaluated.
4.4.2 In vivo toxicity screening
After in vitro cytotoxicity tests, in vivo studies are carried out in order to find out whether the material is irritant. Materials that are known to cause inflammatory reactions are not evaluated with this test. Exposure of the material to the animal may be through touching to skin or injecting or implanting under the skin. The material is applied at least 4 days to study the acute inflammation that results. Histological evaluation is done afterwards.
4.5 Immunological tests
Most biomaterials are composed of natural or artificial polymers. These polymers elicit an immune response after implantation. Such an immune response can create a series of important pathological conditions. For example, the presence of circulating antigens against a particular biomaterial can generate an autoimmune response. The production of immunoglobulin E (IgE) allows for the capability to find out of this allergic phenomenon characterized by immediate hypersensitivity. Such a hypersensitivity response is seen with natural polymers like bovine collagen. T lymphocyte activation represents a necessary step toward acute or chronic rejection of an implant therefore it can be used to verify a material’s immunocompatibility.
4.5.1 Humoral response tests
Antibody production against a particular material must be evaluated by injecting it subcutaneously or intramuscularly into laboratory animals. The material is incorporated into a suspension, if it is not soluble, and preferably mixed with an adjuvant to have a more pronounced effect. Reinjection of antigen 7 days after the first injection is done. After 14 days, when the best antibody production is seen, blood is taken from animal and antibody levels are determined by exposing the blood to antigen and the measuring the amount of antigen-antibody complexes.
4.5.2 Cellular response tests
These tests are designed to detect the response of lymphocytes to soluble substances that are released from the materials during the immune response. These tests usually involve preparing a soluble or colloidal extract of the material to be tested. Then the lymphocytes are cultured in a medium containing a mitogen and the tissue extract. The amount of lymphocyte division is compared with the control groups.
4.6 Studies for the interaction of degradation products with tissues
Radioactively labeled polymers are incubated with tissue homogenates or cells and tissues in culture. The rate of release of radioactivity into the medium is assayed and compared to standards. Release of labeled products into the medium can be used to analyze quantitatively the percentage of the implant that is solubilized. Degradation products can also be tested in cell culture by measuring their effects on cell replication, cloning efficiency, cell viability, cell morphology, and accumulation of intracellular degradation products. Studies are performed using phase-contrast microscopy, labeling techniques, and scanning and transmission electron microscopy.
4.7 Testing for carcinogenic potential
A variety of test are used to detect mutations in the genetic material of a cell induced by a biomaterial. Mutation potential can be studied by exposing the material surface to cultures of mammalian, bacterial, fungal, or yeast cells. Exposure to material results in inhibition of cell replication if a zone of inhibition is observed. Then the amount of inhibition is quantified by the measurement of inhibition zone.
In vivo biocompatibility and biodegradation
of
poly (ethylene carbonate)
M. Dadsetan, E.M. Christenson,
F. Unger, M. Ausborn,
T. Kissel, A. Hiltner, J.M. Anderson
Journal of Controlled Release 93 (2003) 259– 270
5. ıntroductıon
Biodegradable polymers that control drug release through surface erosion are of interest for the delivery of oligopeptides, polypeptides and other macromolecular drugs. These systems may preserve the biological activity of peptides and other labile drugs that are susceptible to hydrolysisPoly(ethylene carbonate) (PEC) is one of the few polymers that undergoes rapid in vivo degradation via surface erosion without undesirable side effects.
Inoue reported that no undesirable
reactions are taking place at the site of implantation of PEC. Acemoğlu et al. found that molecular weight and the carbonate content were identified as important parameters affecting biodegradation of PEC. Stoll et al. hypothesized that PEC degradation was cell mediated and superoxide anion released by adherent cells was involved in in vivo oxidative degradation of PEC. An increased level of superoxide anion production from polymorphonuclear cells (PMNs) in the presence of PEC as well as in vitro testing of PEC degradation in aqueous KO2 at pH 12 confirmed superoxide anion as the source of degradation.
Stoll et al. reported no degradation of PEC with molecular weight of 300–450 kDa in hydrolytic enzymes including lipase, esterase, lysozyme, chymotrypsin, trypsin, papin, pepsin, collagenase, pronase and pronase E. From these results, hydrolytic mechanisms based on hydrolases or aqueous conditions were excluded for biodegradation of PEC.
Although the biodegradation mechanism has been studied extensively, biocompatibility of PEC and PEC degradation products has not been studied in detail. This research has been done to illuminate that point. Biocompatibility was evaluated by monitoring inflammatory leukocyte concentrations in the cage implant exudates. Attenuated total reflectance-Fourier transform infrared (ATR-FTIR) spectroscopy was used to examine the chemical changes of the surface after in vivo degradation.
6. materıals and methods
6.1 Test material
1 x 2.5 cm and 700 μm thick PEC molecules were prepared by molding at 80°C, 10-100 bar pressure for 10 min. Four types of materials were provided; they ranged in molecular weight from 240 to 285 kDa (table 1). The glass transition temperature of PEC was 24–27 °C.
6.2 Cage implantation
The PEC films were weighed and disinfected with an ethanol/water solution (70/30 v/v). One specimen was placed in each cylindrical stainless steel wire mesh cage (3.5 cm length, 1 cm diameter). Each cage was dipped in ethanol/water (70/30 v/v) and rinsed in PBS prior to implantation. The cages were implanted subcutaneously in the back of 3-month-old female Sprague–Dawley rats. Five specimens were implanted for each PEC at each time point. Empty cages were implanted as controls.
6.3 Analysis of in vivo response
Exudate was withdrawn from the cages at 5, 10 and 15 days for determination of total and differential leukocyte concentrations. Rats were sacrificed at 5, 10 and 15 days postimplantation. Total leukocyte counts were performed using a hemacytometer (table 1). One specimen was fixed in glutaraldehyde (2.5%), dried with increasing graded alcohol and gold coated for SEM. One specimen was stained with Wright’s modified stain for optical microscope observation.
6.4 Degradation analysis
The remaining three explanted specimens were sonicated in 1% aqueous Triton X-100 detergent for 10 min, rinsed twice in distilled water and in 70% ethanol solution to remove the cells.
Table 1: Analysis of total and differential in vivo exudate leukocyte concentrations (cells/μl)
The molecular weight of the explanted PECs was determined by gel permeation chromatography (GPC) onto a PSS-SDV linear pre-analytical column system with a multiple angle laser light scattering (MALLS) detector and a RI-detector at ambient temperature. Absolute molecular weights were calculated using the Astra software.
The micro-ATR attachment used a germanium crystal to achieve a sampling depth of approximately 0.1–0.2 μm. ATR-FTIR analysis was performed on film surfaces using a Nicolet 800 FTIR spectrometer.
7. results and dıscussıon
7.1 Biocompatibility
The cage implant system was used to determine the in vivo biocompatibility of PECs with molecular weight 240–285 kDa.
PMNs show the acute inflammatory response at shorter times, whereas increased concentration of macrophages and lymphocytes shows chronic inflammation.
Implantation of the stainless steel cage initiates an inflammatory response. Empty cages, which induce minimal inflammatory response, are used as the control for the implantation procedure. Biocompatibility of the material in the cage is assessed from differences in inflammatory cell concentrations compared to empty cages at different time points.
Table 1 shows the total and differential leukocyte concentrations in the cage implant exudates at 5, 10 and 15 days. Total and differential exudate leukocyte concentrations for all cages containing PEC film and empty cages were maximum at 5 days postimplantation.
The total leukocyte concentration decreased at each subsequent time point. The presence of PEC film did not produce any significant variations in leukocyte concentrations compared to empty cage controls. Thisindicates that PEC and PEC degradation products were biocompatible compared to empty cage controls and induced minimal inflammatory and wound healing responses.
7.2 Cell adhesion and biodegradation
The optical micrographs in Fig. 1 show adherent cells on explanted PEC after Wright’s modified staining. At 5 days postimplantation, primarily macrophages and some small foreign body giant cells (FBGCs) were observed on the PEC surface (Fig.1a). At day 10, most of the cells on PEC film were FBGCs. The FBGCs had increased in size compared to day 5 due to macrophage fusion (Fig. 1b). Pitting was also observed on the PEC surface at this time point. The pits frequently had FBGCs closely adhered to the edges. At day 15, the FBGCs were even larger due to further fusion with other FBGCs and macrophages (Fig. 1c). Pitting covered the entire PEC film by day 15.
The PEC film with adherent cells at 10 days postimplantation is shown in the SEM images in Fig. 2a. The pitting is very obvious. Most of the pits contain FBGCs with their membranes tightly adhered to the edges and surfaces of the pits. Good adhesion of the cell membrane in a pitted region of PEC is seen more clearly in the higher magnification image in Fig. 2b.
It is proposed that degradation occurred at the FBGC/polymer interface where the membrane formed a tight seal with the film surface.
Fig 1. Optical micrographs of Wright’s modified stained cells on explanted PEC1: (a) at 5 days showing primarily adhered macrophages and a few small FBGCs; (b) at 10 days showing larger FBGCs and eroded surface pits of approximately the same size as the FBGCs; (c) at 15 days showing very large FBGCs and surface erosion.
Adhered macrophages and FBGCs can produce reactive oxygen intermediates (ions and free radicals) in response to certain stimuli; these species subsequently initiate degradation at the polymer–cell interface.
The importance of direct cell adhesion to the PEC surface was demonstrated by the observation that; all the PEC samples were completely degraded after subcutaneous implantation for 10 days, whereas PEC in cages experienced an average weight loss of only 14% after 10 days (Stoll et. al. and unpublished data of researchers of this article).
Figure 2. SEM micrographs of adherent cells on the explanted PEC1 surface at 10 days: (a) lower magnification and (b) higher magnification showing intimate contact between a FBGC and an eroded surface pit.
Explanted PEC films after removal of the adherent cells are shown in the SEM images in Fig. 3. Before implantation, the surface of PEC was smooth without any cracking or pitting (Fig. 3a). After 5 days implantation, the film surface contained numerous pits; however, there was no cracking (Fig. 3b). The pits increased in size at day 10 (Fig. 3c). By day 15, the entire surface was eroded to a highly porous texture (Fig. 3d).
Figure 3. SEM micrographs of PEC1 surfaces after removal of adhered cells: (a) before implantation, (b) at 5 days, (c) at 10 days and (d) at 15 days.
Figure 4. SEM micrographs of PEC1 film cross-sections: (a) before implantation, (b) at 5 days, (c) at 10 days and (d) at 15 days.
Cross-sections of the films in Fig. 3 are compared in Fig. 4. After 5 days, the bulk of the film appeared undegraded (Fig. 4b), confirming that degradation at this time point was confined to the surface. After 10 days, a highly porous layer extended from the surface to a depth of approximately 100 μm (Fig. 4c). The solid region in the center appeared to be unaffected, which suggested that degradation proceeded inward by surface erosion. After 15 days, the thickness of the solid core was about 300 Am compared to 700 μm before implantation (Fig. 4d).
The polymer mass after implantation is given in Table 2 together with the film thickness and the molecular weight of remaining PEC. There was almost no mass loss from any of the four materials after 5 days. Thereafter, the remaining polymer mass decreased rapidly. The apparent acceleration in degradation rate with time paralleled the growth in size and coverage of FBGCs seen Fig. 1, and suggested the key role these cells played in the degradation process. The pores of the surface layer were large enough that FBGCs had access to undegraded material far below the film surface.
Data in table 2 is also evident that there was no significant effect of PEC molecular weight on the rate of degradation in the relatively narrow range investigated. This is consistent with results of Acemoglu et al. who reported no effect of molecular weight on biodegradation of PECs in the much broader range of 100–1000 kDa.
7.3 Mechanism of surface erosion
Fig. 5a shows the ATR-FTIR spectrum of PEC film before implantation. Bands at 1741 and 1221 cm-1 are assigned to carbonyl stretching and C-O-C stretching of polycarbonate, respectively.
After 5 days implantation, the carbonate bands at 1741 and 1221 cm-1 decreased in intensity relative to the methylene band at 780 cm-1, and broad new bands appeared at 1654 and 1541 cm-1 (Fig. 5b).
The peak at 1650 cm-1 has been assigned to asymmetric carbonyl stretching of the carbonate ion and the peak at 1540 cm-1 to C-O-C stretching of the carbonate ion.
Compared to the spectrum of 5-day implanted PEC, the new carbonate ion peaks of 15-day implanted PEC at 1654 and 1541 cm-1 were much weaker and the carbonate peaks at 1741 and 1221 cm-1 were stronger. The surface erosion mechanism of degradation can explain the unexpected difference between spectra obtained after 5 days and spectra obtained after longer implantation.
Figure 5. ATR-FTIR spectra of PEC1 film: (a) before implantation, (b) at 5 days and (c) at 15 days. Peaks at 1654 and 1541 cm-1 indicate presence of carbonate ion.
7.4 Action of superoxide ions
It appears that reactive oxygen species, in particular superoxide anion, released from adherent FBGCs on the implanted PEC film, initiate a degradation process that involves formation of carbonate ion.
Random chain scission as the primary degradation mechanism seems unlikely in view of the excellent in vivo biostability of aliphatic polycarbonate soft segments, as compared to polyether soft segments, in polyurethanes.
Two pathways for in vivo degradation of PEC were proposed. In the first, superoxide anion initiates PEC degradation through an oxidation process. In the other, superoxide anion initiates degradation through an alkaline deprotonation mechanism. However, the two possibilities could not be distinguished based on the data presented.
The researchers proposed a deprotonation mechanism that proceeds through a carbonate ion intermediate in accordance with our detection of this species by infrared. Superoxide anion is a strong base, and can deprotonate the PEC hydroxyl end group, leading to an alkoxide ion intermediate (Fig. 7).
Subsequently, the alkoxide ion together with H2O reacts with the carbonate carbonyl and ethylene oxide ion is eliminated through a concerted mechanism of the cyclic intermediate leaving a carbonic acid end group. This is a favorable strategy in enzymatic reactions
The proposed mechanism is an example of neighboring- group participation in which the reaction center interacts with a pair on non-conjugated electrons on the same molecule. Carbonate ion decomposes further to regenerate alkoxide ion, and the reaction sequence repeats to ‘‘unzip’’ the PEC chain.
Figure 7. Superoxide anion initiation of PEC chain degradation.
8. ConclusIon
The finding that implanted cages containing PEC showed no increase in inflammatory cell concentrations compared to empty cage controls indicated that PEC and PEC degradation products are biocompatible and induce minimal inflammatory and wound healing responses.
The PEC film degraded rapidly in vivo losing up to 50% in mass in 15 days.
Observation of FBGCs intimately adhered to the surface at degradation sites and increasing FBGC coverage with time were evidence that reactive species released by FBGCs, in particular superoxide anion, were responsible for chemical degradation of the PEC film.
Recognizing that superoxide anion is a strong base, the researchers suggest that deprotonation of the PEC hydroxyl end group initiates a reaction scheme that leads to unzipping of the PEC chain.
The alkoxide ion together with H2O reacts with carbonate carbonyl to eliminate ethylene oxide through a concerted mechanism of the cyclic intermediate leaving a carbonic acid end group. Further decomposition of carbonate ion regenerates alkoxide ion, and the reaction sequence repeats to ‘‘unzip’’ the PEC chain.
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